## 1 Introduction  
Voice disorders have been estimated to affect approximately 30% of the adult population in the United States at some point in their lives, with up to 7.6% of individuals affected at any given point in time [1; 2]. The most common voice disorders are chronic or recurring conditions believed to result from excessive and/or poorly regulated activity of the perilaryngeal muscles, referred to as vocal hyperfunction. Vocal hyperfunction can be associated with either phonotrauma-induced lesions of the vocal folds (e.g., nodules and polyps) or with dysphonia occurring in the absence of vocal fold trauma or structural or neurological abnormalities (e.g., primary muscle tension dysphonia) [3]. The etiology of phonotraumatic lesions has classically centered on the role of vocal fold impact stress (in the direction of tissue motion at contact) and shear stress (along the tissue surface) during phonation [4]. Insight into the role of collision pressures in the generation of vocal fold lesions is important for any quantification or modeling of phonotrauma since phonotraumatic lesions are widely considered to form due to repetitive stress on the mid-membranous portion of the vocal folds at the location of maximal tissue contact [5; 6; 7]. The direct measurement of vocal fold collision pressure is challenging to carry out in vivo, and only two published studies have attempted to gather data from sensors placed intraglottally during phonation [8; 9]. [PERSON] et al. [9; 10] developed a piezoresistive pressure sensor with a flat frequency response up to 50 kHz and a linear sensing range of 0-14 kPa (0-140 cm H\({}_{2}\)O). The circular pressure-sensing element was 1.8 mm in diameter (0.4 mm thickness) and inserted transorally via a curved cannula. Limited in vivo data were successfully obtained from vocally healthy individuals given challenges related to sensor positioning between two vibrating vocal folds without disrupting phonation and uncertainty as to whether the pressure signal was related to vocal fold collision or acoustic (non-impact) pressures. Vocal fold collision pressures across study participants were reported to be in the range of 4-32 cm H\({}_{2}\)O. However, the criteria by which collision pressure signals in that study were deemed to be of adequate quality were unclear. Only one exemplary waveform exhibited peaks in the pressure signal at the expected vocal fold oscillation rate, but noise in the signal (and a flat-line microphone signal) prevented the ability to confirm whether the peaks were due to impact stress sensing or subglottal (or supraglottal) acoustic pressure sensing. [PERSON] et al. [8] employed a piezoelectric, force-sensitive plate (10 mm \(\times\) 15 mm for a sensing area of 150 mm\({}^{2}\); 0.29 mm thickness) at the end of a transorally positioned rigid cannula that was inserted between the vocal folds. The sensing element exhibited a flat frequency response up to 25 kHz and a linear sensing range for measured forces from a 2.5 mN noise floor up to 200 mN (16-1333 Pa, or 0.16-13.6 cm H\({}_{2}\)O). In that study, peaks in the electroglottography signal and force sensor signal provided an indication of whether the force peaks measured were due to air pressure (force sensor peak lagging behind the electroglottography peak) or vocal fold collision forces (force sensor peak aligned with the electroglottography peak). Intraglottal placement of the force sensor during phonation was considered satisfactory when endoscopic imaging confirmed the desired sensor positioning against one vocal fold, anterior-posterior positioning in the mid-membranous glottis, and inferior-superior positioning of the force sensor in the coronal plane of vocal fold contact (the strike zone). The exemplary waveforms in that paper displayed a force sensor output that resembled a triangular waveform that has not been observed in computational or bench-top models of phonation.