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Airway Gas Flow

Merryn H. Tawhai, Ching‐Long Lin

10.1002/cphy.c100020

Source: Volume 1, Issue 3, July 2011

Published online: July 2011

Full Article on Wiley Online Library



Abstract

Local characteristics of airflow and its global distribution in the lung are determined by interaction between resistance to flow through the airways and the compliance of the tissue, with tissue compliance dominating flow distribution in the healthy lung. Current understanding is that conceptualizing the airways of the lung as a system of smooth adjoined cylinders through which air traverses laminarly is insufficient for understanding flow and energy dissipation and is particularly poor for predicting physiologically realistic transport of particles by the airflow. With rapid advances in medical imaging, computer technologies, and computational techniques, computational fluid dynamics is now becoming a viable tool for providing detailed information on the mechanics of airflow in the human respiratory tract. Studies using such techniques have shown that the upper airway (specifically its development of a turbulent laryngeal jet in the trachea), airway geometry, branching and rotation angle, and the pattern of joining of successive bifurcations are important in determining airflow structures. It is now possible to compute airflow in physical domains that are anatomically accurate and subject specific, enabling comparisons among intersubjects, that among subjects of different ages, and that among different species. © 2011 American Physiological Society. Compr Physiol 1:1135‐1157, 2011.

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Figure 1. Figure 1.

Flow regimes of the conducting airway categorized on the basis of a dimensionless frequency α2 (where α is the Womersley number) and a dimensionless stroke length L/a. The (I) unsteady, (II) viscous, and (IIIa, IIIb) convective flow regimes are classified according to Jan et al. 44.
Figure 2. Figure 2.

Coherent vortical structures found in a shear‐driven flat‐plate turbulent boundary layer. From Head and Bandyopadhyay 38, see also Robinson 78.
Figure 3. Figure 3.

A visualization of a plane mixing layer between helium (upper) at a speed of 10.1 m/s and nitrogen (lower) at 3.8 m/s. From 10.
Figure 4. Figure 4.

Contours of streamwise velocity in pipe flow through a 50% constriction (D, diameter of the tube) at Re = 2000.
Figure 5. Figure 5.

Velocity vector field and contours of speed at cross sections within the trachea (A and D), the left main bronchus (B and E), and the right main bronchus (C and F) for computational fluid dynamics simulation of airflow with (case 1: A‐C) and without (case 2: D‐F) including the upper airways. Reproduced with permission from Lin et al. 57.
Figure 6. Figure 6.

Outlet velocity vectors (pink) and pressure contours for three different downstream boundary conditions: (A) image‐based boundary condition; (B) uniform velocity; (C) uniform pressure. The pressure at the trachea pt is used as a reference. (D) Flow partitions in the five lobes. LUL, left upper lobe; LLL, left lower lobe; RUL, right upper lobe; RML, right middle lobe; RLL, right lower lobe. Reproduced with permission from Yin et al. 111.
Figure 7. Figure 7.

Large eddy simulation–generated airflow structures in two airway models based on multidetector row computed tomography. Laryngeal jet structures are indicated by isosurfaces of mean speed and contours in a vertical plane cutting through the vocal cord and the trachea. Reproduced with permission from Choi et al. 16.
Figure 8. Figure 8.

Velocity distributions in the bifurcation plane and two cross sections at Re = 740, L/a = 15, α = 7 (point B in Figure 1) at: (A) end inspiration and (B) end expiration. A sinusoidal waveform is imposed at the inlet of the parent branch with 0 ≤ t/T ≤ 0.5 (0.5 ≤ t/T ≤ 1) for inspiratory (expiratory) phase. Blue, negative axial velocity to the left; red, positive axial velocity to the right. The vector length in the two cross sections is magnified four times for clarity. Reproduced with permission from Choi et al. 17.
Figure 9. Figure 9.

Velocity vectors at end inspiration (t/T = 0.47) and end expiration (t/T = 0.97) in the bifurcation plane at the bifurcation between generations 2 and 3. Red (blue), positive (negative) axial velocity in the parent branch to the right (left). (A and B) NORM (a normal breathing case); (C and D) HFNR (a high‐frequency normal Re case); and (E and F) HFOV (a high‐frequency oscillatory ventilation case). Re at the parent branch is 64 for (A‐D) and 182 for (E and F). Reproduced with permission from Choi et al. 17.
Figure 10. Figure 10.

Instantaneous particle transport profiles and contours of velocity magnitude for: (A) 2.5‐μm particles, (B) 20‐μm particles.
Figure 11. Figure 11.

Time histories of the maximum shear stresses in a rigid airway, a flexible airway, and a flexible airway wall with parenchymal tethering. Reproduced with permission from Xia et al. 109.
Figure 12. Figure 12.

Wall shear stress distribution at t = 0.25T, Re = 183 in (A) rigid airway; (B) flexible airway; and (C) flexible airway wall with parenchymal tethering. Wall shear stress is shown in units of Pa, with red indicating highest stress and dark blue lowest stress. Reproduced with permission from Xia et al. 109.
Figure 13. Figure 13.

An illustration of registration‐derived moving airway geometries at four different time points in a breathing cycle: (A), t = 0; (B), t = T/6; (C), t = T/3; and (D), t = T/2, with T as the period. This moving model is derived from a pair of volumetric multidetector row computed tomographic data sets with a lung volume change of 1.5 liters and the geometries at (A) and (D) correspond to the lower and higher lung volumes, respectively.
Figure 14. Figure 14.

Three‐ and one‐dimensional coupled airway tree and five lobes at (A) the minimum volume (55% vital capacity [VC]) and (B) the maximum lung volume (85%VC). (C) The distribution of deformation as a function of normalized lung height from the apex (100%) of the lung to the base (0%) calculated from the Jacobian‐estimated volume change. ΔV and ΔL denote volume and length changes. The subscript “max” corresponds to the maximum lung volume.


Figure 1.

Flow regimes of the conducting airway categorized on the basis of a dimensionless frequency α2 (where α is the Womersley number) and a dimensionless stroke length L/a. The (I) unsteady, (II) viscous, and (IIIa, IIIb) convective flow regimes are classified according to Jan et al. 44.



Figure 2.

Coherent vortical structures found in a shear‐driven flat‐plate turbulent boundary layer. From Head and Bandyopadhyay 38, see also Robinson 78.



Figure 3.

A visualization of a plane mixing layer between helium (upper) at a speed of 10.1 m/s and nitrogen (lower) at 3.8 m/s. From 10.



Figure 4.

Contours of streamwise velocity in pipe flow through a 50% constriction (D, diameter of the tube) at Re = 2000.



Figure 5.

Velocity vector field and contours of speed at cross sections within the trachea (A and D), the left main bronchus (B and E), and the right main bronchus (C and F) for computational fluid dynamics simulation of airflow with (case 1: A‐C) and without (case 2: D‐F) including the upper airways. Reproduced with permission from Lin et al. 57.



Figure 6.

Outlet velocity vectors (pink) and pressure contours for three different downstream boundary conditions: (A) image‐based boundary condition; (B) uniform velocity; (C) uniform pressure. The pressure at the trachea pt is used as a reference. (D) Flow partitions in the five lobes. LUL, left upper lobe; LLL, left lower lobe; RUL, right upper lobe; RML, right middle lobe; RLL, right lower lobe. Reproduced with permission from Yin et al. 111.



Figure 7.

Large eddy simulation–generated airflow structures in two airway models based on multidetector row computed tomography. Laryngeal jet structures are indicated by isosurfaces of mean speed and contours in a vertical plane cutting through the vocal cord and the trachea. Reproduced with permission from Choi et al. 16.



Figure 8.

Velocity distributions in the bifurcation plane and two cross sections at Re = 740, L/a = 15, α = 7 (point B in Figure 1) at: (A) end inspiration and (B) end expiration. A sinusoidal waveform is imposed at the inlet of the parent branch with 0 ≤ t/T ≤ 0.5 (0.5 ≤ t/T ≤ 1) for inspiratory (expiratory) phase. Blue, negative axial velocity to the left; red, positive axial velocity to the right. The vector length in the two cross sections is magnified four times for clarity. Reproduced with permission from Choi et al. 17.



Figure 9.

Velocity vectors at end inspiration (t/T = 0.47) and end expiration (t/T = 0.97) in the bifurcation plane at the bifurcation between generations 2 and 3. Red (blue), positive (negative) axial velocity in the parent branch to the right (left). (A and B) NORM (a normal breathing case); (C and D) HFNR (a high‐frequency normal Re case); and (E and F) HFOV (a high‐frequency oscillatory ventilation case). Re at the parent branch is 64 for (A‐D) and 182 for (E and F). Reproduced with permission from Choi et al. 17.



Figure 10.

Instantaneous particle transport profiles and contours of velocity magnitude for: (A) 2.5‐μm particles, (B) 20‐μm particles.



Figure 11.

Time histories of the maximum shear stresses in a rigid airway, a flexible airway, and a flexible airway wall with parenchymal tethering. Reproduced with permission from Xia et al. 109.



Figure 12.

Wall shear stress distribution at t = 0.25T, Re = 183 in (A) rigid airway; (B) flexible airway; and (C) flexible airway wall with parenchymal tethering. Wall shear stress is shown in units of Pa, with red indicating highest stress and dark blue lowest stress. Reproduced with permission from Xia et al. 109.



Figure 13.

An illustration of registration‐derived moving airway geometries at four different time points in a breathing cycle: (A), t = 0; (B), t = T/6; (C), t = T/3; and (D), t = T/2, with T as the period. This moving model is derived from a pair of volumetric multidetector row computed tomographic data sets with a lung volume change of 1.5 liters and the geometries at (A) and (D) correspond to the lower and higher lung volumes, respectively.



Figure 14.

Three‐ and one‐dimensional coupled airway tree and five lobes at (A) the minimum volume (55% vital capacity [VC]) and (B) the maximum lung volume (85%VC). (C) The distribution of deformation as a function of normalized lung height from the apex (100%) of the lung to the base (0%) calculated from the Jacobian‐estimated volume change. ΔV and ΔL denote volume and length changes. The subscript “max” corresponds to the maximum lung volume.

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Article Title: Airway Gas Flow
 

How to Cite

Merryn H. Tawhai, Ching‐Long Lin. Airway Gas Flow. Compr Physiol 2011, 1: 1135-1157. doi: 10.1002/cphy.c100020