## 1 Introduction  
Voiced speech is produced as the vocal folds (VFs) undergo periodic oscillations arising from fluid-structure-acoustic [1; 2; 3; 4] interactions. Each phonatory cycle is characterized by the opening and closing of the glottal area between the VFs [5], where voice quality is often dependent upon the degree of glottal closure [6]. Excessive contact forces that occur during glottal closure may, however, give rise to phonotrauma, which can be acute, or chronic [7]. Chronic behaviors can lead to phonotraumatic vocal hyperfunction [8], where compensation for a pathology, such as nodules, further exacerbates the underlying problem leading to a deleterious cycle of continual and increasing vocal fold tissue damage. Understanding physiological variables that influence VF contact forces can provide unique insightsinto the diagnosis, treatment, and remediation of voice disorders associated with phonotraumatic vocal hyperfunction. Both direct and indirect measures of contact forces have been investigated using theoretical [9; 10], experimental [11; 12; 13; 14], and computational approaches [15; 16; 17; 18; 19; 20; 21]. Efforts focused on directly measuring VF contact are of particular interest as successful techniques would be easily applicable to clinical investigations. A representative intraglottal pressure waveform from early efforts investigating contact pressure in excised canine larynges is presented in Figure 1[11], where two clear peaks were identified in the time history of the pressure, with the primary peak arising from VF contact. As the VFs began to open in the Pre-Open Phase the wall-mounted pressure sensor was exposed to the subglottal pressure while the VFs were still closed, leading to the pressure rising before reaching a secondary peak during the Open Phase. The magnitude of the maximum contact pressure was correlated with the subglottal pressure, reaching values that were as high, or higher, than the subglottal pressure [11]. This cyclical, high amplitude, impact pressure is believed to produce cumulative damage in the VF structure, as most trauma becomes evident after prolonged voice usage [22; 23; 24]. Direct _in vivo_ contact pressure measurements have, however, remained largely elusive due to the practical difficulties associated with sensor size, placement, and accuracy. Challenges arise due to (1) physical intrusion of the sensor, (2) the size of the sensor relative to the contact area, which may bias the recorded values due to poor spatial resolution, and (3) accurate positioning of any sensing element within the region of VF contact, as the VF contact length in the inferior-superior direction is \(\mathcal{O}(\mathrm{mm})\) as the VFs are vibrating at \(\mathcal{O}(100\,\mathrm{Hz})\). [PERSON] and [PERSON], (1994) [11] acknowledged that the amount the pressure transducer was recessed into, or protruded above, the surrounding surface against which the VFs impacted had a significant influence on the magnitude of the measured contact pressure. While they chose to position the sensor such that it protruded slightly (\(\sim 0.5\,\mathrm{mm}\)) above the contact wall, no validation with a ground-truth measure of pressure was performed. Similar challenges were observed in the work of [PERSON] et. al., (2005) [25], who designed a custom, thin-film piezoelectric force-sensing element. Unfortunately, the size (\(\sim\)\(10\,\mathrm{mm}\times 15\,\mathrm{mm}\) and \(0.29\,\mathrm{mm}\) thick) was on order with the size of the glottal opening. Additional approaches have sought to combat this challenge through implementation of flush-mounted sensors with synthetic, silicone VF models oscillating in a hemilaryngeal orientation [14; 26]. To increase the spatial resolution of contact pressure measurements, smaller sensors have been utilized [14; 27; 28]. The first _in vivo_ VF contact pressure measurements were carried out using a \(1.8\,\mathrm{mm}\) diameter piezoresistive sensor mounted on a curved, plastic-insulated metal cannula [28].