## 1 Introduction  
The etiology of trauma-induced vocal fold (VF) lesions, such as nodules and polyps, has been long associated with detrimental vocal patterns and compensatory behaviors that result in increased VF collision pressures [1; 2; 3; 4; 5]. Unfortunately, a direct in vivo quantification of VF collision pressure is not feasible in routine clinical examinations due to the invasive nature of the procedure, i.e., placing a pressure probe in the glottis constitutes a risk for the VF tissue [6; 7; 8]. Therefore, current objective assessments of VF collision in clinical environments relies on indirect methods (e.g., aerodynamic [4], acoustic [9], ambulatory [10], etc.) and does not take advantage of the most direct and common procedure to clinically assess VF function: Laryngeal endoscopy. At the same time, the recent development of laryngeal high-speed videoendoscopy (HSV), with frame rates above 4000 fps, has significantly improved the temporal resolution of this medical examination, thus allowing for an accurate visualization and quantification of vocal fold vibration patterns that are challenging to capture using standard stroboscopic imaging methods [11; 12; 13; 14]. In spite of these benefits, the adoption of HSV into the clinical practice has been slow, in part because current analysis methods and tools to navigate the data are not yet sufficiently intuitive or useful to the clinician [11; 12]. Thus, there is a clear need for enhanced analysis methods for laryngeal endoscopy, and given the high prevalence of nodules and polyps [15], the estimation of VF collision pressure directly from endoscopic observations is poised to have a strong impact in the clinical practice. In spite of the significant interest for directly assessing VF collision, only a few studies have been able to gather contact pressure data in human subjects by placing pressure probes at the glottis [6; 7; 8]. Sensing contact pressure is a challenge in terms of probe size (flow and tissue interference), sensor bandwidth, risk to the VF tissue, and subject tolerance during the examination. Nevertheless, these pilot studies reported values for peak collision pressure of 0.5 to 3.0 kPa [7] in normal participants and 1 to 4 kPa in subjects with organic pathologies, wherein the values were larger near the lesions [6]. A stronger collision has been associated with an increased voice intensity [8]. These results are consistent with intraglottal pressures reported from excised canine hemilarynges [16; 17; 18], where it was observed that the impact produced sharp pressure pulses that were positively correlated with subglottal pressure and VF elongation. In addition to clinical measures, numerical approaches have been used to investigate the role of collision pressures in the development of VF lesions. Using internal stress or ad hoc solutions for computing contact, finite element studies have shown that (1) the elastic forces within the VF tissue are responsible for the impact strength [19]; (2) there is a relationship between the subglottal pressure and peak collision [19; 20]; and (3) VF impact causes a transient in the tissue stress that can be related to fatigue damage [20] and a reduction of hydration [21]. More recently, models of collision have incorporated descriptions of fluid displacement in the VF cover [22] and asymmetric conditions [23] that better represent the underlying physics; these studies yielded VF contact pressures in the same range as the experimental data. Silicone VF replicas have also been used to study VF collision [24; 25]. In these studies, a digital image correlation (DIC) approach and a Hertzian impact model [26; 27] were utilized to quantify surface strain distributions on the superior VF surface and to obtain collision pressures in silicone models of the VFs. The Hertzian contact model has also been used to quantify collision forces in numerically reduced order models of the vocal folds [28; 29].