Several factors besides the elastic recoil of the lungs and the chest wall must be overcome to move air into or out of the lungs. These factors include the inertia of the respiratory system, the frictional resistance of the lung and chest wall tissue, and the frictional resistance of the airways to the flow of air. The inertia of the system is negligible. Pulmonary tissue resistance is caused by the friction encountered as the lung tissues move against each other or the chest wall as the lung expands. The airways resistance plus the pulmonary tissue resistance is often referred to as the pulmonary resistance. Pulmonary tissue resistance normally contributes about 20% of the pulmonary resistance, with airways resistance responsible for the other 80%. Pulmonary tissue resistance can be increased in such conditions as pulmonary sarcoidosis, silicosis, asbestosis, and fibrosis. Because airways resistance is the major component of the total resistance and because it can increase tremendously both in healthy people and in those suffering from various diseases, the remainder of this chapter will concentrate on airways resistance.
Laminar, Turbulent, and Transitional Flow
Generally, the relationship among pressure, flow, and resistance is stated as
This means that resistance is a meaningful term only during flow. When airflow is considered, the units of resistance are usually cm H2O/L/s.
The resistance to airflow is analogous to electrical resistance in that resistances in series are added directly:
Resistances in parallel are added as reciprocals:
Understanding and quantifying the resistance to airflow in the conducting system of the lungs is difficult because of the nature of the airways themselves. It is relatively easy to inspect the resistance to airflow in a single, unbranched, indistensible tube; however, the ever-branching, narrowing, distensible, and compressible system of airways complicates the analysis of the factors contributing to airways resistance. Therefore, equations can only approximate clinical situations.
Airflow, like that of other fluids, can occur as either laminar or turbulent flow.
As seen in Figure 2–16, laminar flow (or streamline flow) consists of a number of concentrically arranged cylinders of air flowing at different rates. This telescope like arrangement is such that the cylinder closest to the wall of the vessel has the slowest velocity because of frictional forces with the wall; the pathway in the center of the vessel has the highest velocity.
Illustration of laminar, turbulent, and transitional airflow.
When a fluid such as air flows through rigid, smooth-bore tubes, its behavior is governed by Poiseuille’s law. The pressure difference is directly proportional to the flow times the resistance if flow is laminar:
where ΔP is the pressure difference,
is the airflow, and R1
is the resistance.
According to Poiseuille’s law, the resistance is directly proportional to the viscosity of the fluid and the length of the tube and is inversely proportional to the fourth power of the radius of the tube:
where η is the viscosity of fluid, l is the length of tube, and r is the radius of tube.
Note that if the radius is cut in half, the resistance is multiplied by 16 because the resistance is inversely proportional to the radius to the fourth power.
Flow changes from laminar to turbulent when Reynold’s number exceeds 2000. Reynold’s number is a dimensionless number equal to the density of the fluid times the velocity of the fluid times the diameter of the tube divided by the viscosity of the fluid:
where ρ is the density of fluid, Ve is the linear velocity of fluid, D is the diameter of tube, and η is the viscosity of fluid.
During turbulent flow, the relationship among the pressure difference, flow, and resistance changes. Because the pressure difference is proportional to the flow squared, much greater pressure differences are required to generate the same airflow. The resistance term is influenced more by the density than it is by the viscosity during turbulent flow:
Transitional flow is a mixture of laminar and turbulent flow. This type of flow often occurs at branch points or points distal to partial obstructions.
Turbulent flow tends to occur if airflow is high, gas density is high, the tube radius is large, or all 3 conditions exist. During turbulent flow, flow is inversely proportional to gas density, but viscosity is unimportant as the concentric cylinders of flow (the lamina) break down. True laminar flow probably occurs only in the smallest airways, where the linear velocity of airflow is extremely low. Linear velocity (cm/s) is equal to the flow (cm3/s) divided by the cross-sectional area. The total cross-sectional area of the smallest airways is very large (see Chapter 1), and so the linear velocity of airflow is very low. The airflow in the trachea and larger airways is usually either turbulent or transitional.
Distribution of Airways Resistance
In a normal adult about 35% to 50% of the total resistance to airflow is located in the upper airways: the nose, nasal turbinates, oropharynx, nasopharynx, and larynx. Resistance is higher when one breathes through the nose than when one breathes through the mouth.
The vocal cords open slightly during normal inspirations and close slightly during expirations. During deep inspirations, they open widely. The muscles of the oropharynx also contract during normal inspirations, which dilates and stabilizes the upper airway. During deep forced inspirations, the development of negative pressure could cause the upper airway to be pulled inward and partly or completely obstruct airflow. Reflex contraction of these pharyngeal dilator muscles normally keeps the airway open (see Figure 2–25).
As for the tracheobronchial tree, the component with the highest individual resistance is the smallest airway, which has the smallest radius. Nevertheless, because the smallest airways are arranged in parallel, their resistances add as reciprocals, so that the total resistance to airflow offered by the numerous small airways is extremely low during normal, quiet breathing. Therefore, under normal circumstances the greatest resistance to airflow resides in the large to medium-sized bronchi.
Control of Bronchial Smooth Muscle
The smooth muscle of the airways from the trachea down to the alveolar ducts is under the control of efferent fibers of the autonomic nervous system. Stimulation of the cholinergic parasympathetic postganglionic fibers causes constriction of bronchial smooth muscle as well as increased glandular mucus secretion. The preganglionic fibers travel in the vagus. Stimulation of the adrenergic sympathetic fibers causes dilation of bronchial and bronchiolar smooth muscle as well as inhibition of glandular secretion. This dilation of the airways smooth muscle is mediated by beta2 (β2) receptors, which predominate in the airways. Selective stimulation of the alpha (α) receptors with pharmacologic agents causes bronchoconstriction. Adrenergic transmitters carried in the blood may be as important as those released from the sympathetic nerves in causing bronchodilation. The bronchial smooth muscle is normally under greater parasympathetic than sympathetic tone.
Inhalation of chemical irritants, smoke, or dust; stimulation of the arterial chemoreceptors; and substances such as histamine cause reflex constriction of the airways. Decreased CO2 in the branches of the conducting system causes a local constriction of the smooth muscle of the nearby airways; increased CO2 or decreased O2 causes a local dilation. This may help balance ventilation and perfusion (see Chapter 5). Many other substances can have direct or indirect effects on airway smooth muscle (Table 2–2). Leukotrienes usually cause bronchoconstriction, as do some prostaglandins.
Table 2–2. Active Control of the Airways ||Download (.pdf)
Table 2–2. Active Control of the Airways
- Parasympathetic stimulation
- Thromboxane A2
- α-Adrenergic agonists
- Decreased in small airways
- Sympathetic stimulation (β2 receptors)
- Circulating β2 agonists
- Nitric oxide
- Increased in small airways
- Decreased in small airways
Lung Volume and Airways Resistance
Airways resistance decreases with increasing lung volume, as shown in Figure 2–17 (normal curve). This relationship is still present in an emphysematous lung (“Abnormal” in Figure 2–17), although in emphysema the resistance is higher than that in a healthy lung, especially at low lung volumes.
Relationship between lung volume and airways resistance. Total lung capacity is at right; residual volume is at left. Solid line = normal lung; dashed line = abnormal (emphysematous) lung. (Reproduced with permission from Murray, 1972.)
There are 2 reasons for this relationship; both mainly involve the small airways, which, as described in Chapter 1, have little or no cartilaginous support. The small airways are therefore rather distensible and also compressible. Thus, the transmural pressure difference across the wall of the small airways is an important determinant of the radius of the small airways: Since resistance is inversely proportional to the radius to the fourth power, changes in the radii of small airways can cause dramatic changes in airways resistance, even with so many parallel pathways. To increase lung volume, a person breathing normally takes a “deep breath,” that is, makes a strong inspiratory effort. This effort causes intrapleural pressure to become much more negative than the –7 or –10 cm H2O seen in a normal, quiet breath. The transmural pressure difference across the wall becomes much more positive, and small airways are distended.
A second reason for the decreased airways resistance seen at higher lung volumes is that the so-called traction on the small airways increases. As shown in the schematic drawing in Figure 2–18 (see also the alveolar duct in Figure 1–3), the small airways traveling through the lung form attachments to the walls of alveoli. As the alveoli expand during the course of a deep inspiration, the elastic recoil in their walls increases; this elastic recoil is transmitted to the attachments at the airway, pulling it open.
Representation of “traction” of the alveolar septa on a small distensible airway. A: Cross-sectional view. Compare this figure with the picture of the alveolar duct in Figure 1–3. B: View along the length of the small airway. Compare this figure with Figure 2–2A.
Dynamic Compression of Airways
Airways resistance is extremely high at low lung volumes, as can be seen in Figure 2–17. To achieve low lung volumes, a person must make a forced expiratory effort by contracting the muscles of expiration, mainly the abdominal and internal intercostal muscles. This effort generates positive intrapleural pressure, which can be as high as 120 cm H2O during a maximal forced expiratory effort. (Maximal inspiratory intrapleural pressures can be as low as –80 cm H2O.)
The effect of this high positive intrapleural pressure on the transmural pressure gradient during a forced expiration can be seen at right in Figure 2–19, a schematic drawing of a single alveolus and airway.
Schematic diagram illustrating dynamic compression of airways and the equal pressure point hypothesis during a forced expiration. Left: Passive (eupneic) expiration. Intrapleural pressure is –8 cm H2O, alveolar elastic recoil pressure is +10 cm H2O, and alveolar pressure is +2 cm H2O. Right: Forced expiration at the same lung volume. Intrapleural pressure is +25 cm H2O, alveolar elastic recoil pressure is +10 cm H2O, and alveolar pressure is +35 cm H2O.
At this instant, during the course of a forced expiration, the muscles of expiration are generating a positive intrapleural pressure of +25 cm H2O. Pressure in the alveolus is greater than intrapleural pressure because of the alveolar elastic recoil pressure of +10 cm H2O, which together with intrapleural pressure, gives an alveolar pressure of +35 cm H2O. The alveolar elastic recoil pressure decreases at lower lung volumes because the alveolus is not as distended. In the figure, a gradient has been established from the alveolar pressure of +35 cm H2O to the atmospheric pressure of 0 cm H2O. If the airways were rigid and incompressible, the large expiratory pressure gradient would generate very high rates of airflow.
However, the airways are not uniformly rigid and the smallest airways, which have no cartilaginous support and rely on the traction of alveolar septa to help keep them open, may be compressed or may even collapse. Whether or not they actually collapse depends on the transmural pressure gradient across the walls of the smallest airways. Small airway collapse is the main reason that airways resistance appears to be approaching infinity at low lung volumes in Figure 2–17
The situation during a normal passive expiration at the same lung volume (note the same alveolar elastic recoil pressure) is shown in the left part of Figure 2–19. The transmural pressure gradient across the smallest airways is
tending to hold the airway open. During the forced expiration at right, the transmural pressure gradient is 30 cm H2O − 25 cm H2O, or only 5 cm H2O holding the airway open. The airway may then be slightly compressed, and its resistance to airflow will be even greater than during the passive expiration. This increased resistance during a forced expiration is called dynamic compression of airways.
Consider what must occur during a maximal forced expiration. As the expiratory effort is increased to attain a lower and lower lung volume, intrapleural pressure is getting more and more positive, and more and more dynamic compression will occur. Furthermore, as lung volume decreases, there will be less alveolar elastic recoil pressure and the difference between alveolar pressure and intrapleural pressure will decrease.
One way of looking at this process is the equal pressure point hypothesis. (Another explanation of flow limitation during forced expiration, the wave speed flow-limiting mechanism, is too complex to discuss here.) At any instant during a forced expiration, there is a point along the airways where the pressure inside the airway is just equal to the pressure outside the airway. At that point the transmural pressure gradient is 0 (note the arrows in Figure 2–19). Above that point, the transmural pressure gradient is negative: The pressure outside the airway is greater than the pressure inside it, and the airway will collapse if cartilaginous support or alveolar septal traction is insufficient to keep it open.
As the forced expiratory effort continues, the equal pressure point is likely to move down the airway toward the alveoli from larger to smaller airways. This occurs because, as the muscular effort increases, intrapleural pressure increases and because, as lung volume decreases, alveolar elastic recoil pressure decreases. As the equal pressure point moves down the airway, dynamic compression increases and the airways ultimately begin to collapse. This airways closure can be demonstrated only at especially low lung volumes in healthy subjects, but the closing volume may occur at higher lung volumes in patients with emphysema, as will be discussed at the end of this chapter. (Note that the point at which airways resistance approaches infinity occurs at a much higher volume in the “abnormal” lungs in Figure 2–17.) The closing volume test itself will be discussed in Chapter 3.
Consider the pressure gradient for airflow during a forced expiration. During a passive expiration the pressure gradient for airflow
is simply alveolar pressure minus atmospheric pressure. But if dynamic compression occurs, the effective pressure gradient is alveolar pressure minus intrapleural pressure
(which equals the alveolar elastic recoil pressure) because intrapleural pressure is greater than atmospheric pressure and because intrapleural pressure can exert its effects on the compressible portion of the airways.
Thus, during a forced expiration, when intrapleural pressure becomes positive and dynamic compression occurs, the effective driving pressure for airflow from the lung is the alveolar elastic recoil pressure. Alveolar elastic recoil is also important in opposing dynamic compression of the airways because of its role in the traction of the alveolar septa on small airways, as shown in Figure 2–18. The effects of alveolar elastic recoil on airflow during a forced expiration are illustrated in Figure 2–20.
Representation of the effects of alveolar elastic recoil on airflow during a forced expiration. When dynamic compression occurs, alveolar elastic recoil helps to oppose it by traction on the small airways. The alveolar elastic recoil pressure becomes the effective driving pressure for airflow from the lung. PA = alveolar pressure; Ppl = intrapleural pressure; Pel = alveolar elastic recoil pressure.
The Bernoulli principle may also play a role in dynamic compression of the airways. For an ideal fluid with no viscosity, as the linear velocity of the fluid flow increases, the pressure exerted by the fluid on the walls of the vessel (the “lateral pressure”) decreases. Therefore, as the velocity of airflow in the small compressible airways increases during a forced expiration, the pressure inside the vessel decreases. This could contribute to a decreased or more negative transmural pressure difference across the vessel wall.
Assessment of Airways Resistance
The resistance to airflow cannot be measured directly but must be calculated from the pressure difference and airflow during a breath:
This formula is an approximation because it presumes that all airflow is laminar, which is not true. But there is a second problem: How can the pressure gradient be determined? To know the pressure gradient, the alveolar pressure—which also cannot be measured directly—must be known. Alveolar pressure can be calculated using a body plethysmograph, an expensive piece of equipment described in detail in the next chapter, but this procedure is not often done. Instead, airways resistance is usually assessed indirectly. The assessment of airways resistance during expiration will be emphasized because it is of interest in patients with emphysema, chronic bronchitis, and asthma.
One way of assessing expiratory airways resistance is to look at the results of a forced expiration into a spirometer, as shown in Figure 2–21. This measurement is called a forced vital capacity (FVC). The VC is the volume of air a subject is able to expire after a maximal inspiration to the total lung capacity (TLC). An FVC means that a maximal expiratory effort was made during this maneuver.
Forced vital capacity (FVC maneuver using a water-seal spirometer). (See Figure 3–4 for a diagram of a spirometer.) Upper trace: FVC from a normal subject. Lower trace: FVC from a patient with obstructive disease. FEV1 = forced expiratory volume in the first second; FEF25%–75% = forced expiratory flow between 25% and 75% of the FVC. Bottom traces: Similar curves obtained from a more commonly used rolling seal spirometer. Note that the total lung capacity (TLC) is at the bottom of the curves and the residual volumes (RVs) are at the top; volume therefore refers to the volume exhaled into the spirometer in the bottom trace. The time scale is from left to right.
In an FVC test, a person makes a maximal inspiration to the TLC. After a moment, he or she makes a maximal forced expiratory effort, blowing as much air as possible out of the lungs. At this point, only a residual volume (RV) of air is left in the lungs. (The lung volumes will be described in detail in the next chapter.) This procedure takes only a few seconds, as can be seen on the time scale.
The part of the curve most sensitive to changes in expiratory airways resistance is the first second of expiration. The volume of air expired in the first second of expiration (the FEV1, or forced expiratory volume in 1 second), especially when expressed as a ratio with the total amount of air expired during the FVC, is a good index of expiratory airways resistance. In normal young subjects, the FEV1/FVC is greater than 0.80; that is, at least 80% of the FVC is expired in the first second. An FEV1/FVC of 75% would be more likely in an older person. A patient with airway obstruction caused by an episode of asthma, for example, would be expected to have an FEV1/FVC far below 0.80, as shown in the middle and bottom panels in Figure 2–21.
The bottom panel of Figure 2–21 shows similar FVC curves that would be obtained from a commonly used rolling seal spirometer. The curves are reversed right to left and upside down if they are compared with those in the top and middle panels. The TLC is at the bottom left, and the RVs are at the top right. The time scale is left to right. Note the calculations of the FEV1 to FVC ratios.
Another way of expressing the same information is the FEF25%–75%, or forced (mid) expiratory flow rate (formerly called the MMFR, or maximal midexpiratory flow rate). This variable is simply the slope of a line drawn between the points on the expiratory curve at 25% and 75% of the FVC. In cases of airway obstruction, this line is not nearly as steep as it is on a curve obtained from someone with normal airways resistance. The FEV1/FVC is usually considered to represent larger airways, the FEF25%–75%, smaller to medium-sized airways.
The main concept underlying these pulmonary function tests is that elevated airways resistance takes time to overcome.
Isovolumetric Pressure-Flow Curve
The isovolumetric pressure-flow technique is not used clinically because intrapleural pressure must be determined and the data obtained are tedious to plot. Analysis of the results obtained from this test, however, demonstrates several points we have already discussed. Isovolumetric pressure-flow curves are obtained by having a subject make repeated expiratory maneuvers with different degrees of effort. Intrapleural pressures are determined with an esophageal balloon, lung volumes are determined with a spirometer, and airflow rates are determined by using a pneumotachograph. The pressure-flow relationship for each of the expiratory maneuvers of various efforts is plotted on a curve for a particular lung volume. With each expiratory effort, as the lung volume passes through the chosen volume, the intrapleural pressure (approximating the expiratory effort) is plotted against the expiratory flow achieved. For example, the middle curve of Figure 2–22 was constructed by determining the intrapleural pressure and airflow for each expiratory maneuver as the subject’s lung volume passed through 50% of the VC. Therefore, none of the 3 curves in Figure 2–22 is really a continuous line; each curve is constructed from individual data points.
Isovolumetric pressure-flow curves at 3 different lung volumes: 75%, 50%, and 25% of the vital capacity (VC). (Reproduced with permission from Hyatt, 1965.)
The middle curve in Figure 2–22 demonstrates dynamic compression and supports the equal pressure point hypothesis. At this lung volume, at which elastic recoil of the alveoli should be the same no matter what the expiratory effort, with increasing expiratory effort airflow increases up to a point. Beyond that point, generating more positive intrapleural pressure does not increase airflow: It becomes effort-independent. Airways resistance must be increasing with increasing expiratory effort. Airflow has become independent of effort because of greater dynamic compression with more positive intrapleural pressures. The equal pressure point has moved to compressible small airways and is fixed there. Note that at even lower lung volumes (25% of the VC), at which there is less alveolar elastic recoil to provide traction on small airways, this occurs with lower maximal airflow rates. In other words, because alveolar pressure equals the sum of the intrapleural pressure and the alveolar elastic recoil pressure during a forced expiration at a given lung volume, the driving pressure for airflow becomes independent of expiratory muscle effort because increasing the intrapleural pressure increases the alveolar pressure by the same amount. Only the alveolar elastic recoil, which is constant at a given lung volume, drives air out of the lung. At high lung volumes (75% of VC), airflow increases steadily with increasing effort. It is entirely effort-dependent because alveolar elastic recoil pressure is high (which increases both the alveolar septal traction on small airways and the pressure gradient for airflow) and because highly positive intrapleural pressures cannot be attained at such high lung volumes with the airway wide open.
These same principles are demonstrated in the expiratory portion of flow-volume curves (Figure 2–23).
Flow-volume curves of varying intensities, demonstrating effort dependence at high lung volumes and effort independence at low lung volumes. Note that there is normally no effort independence in inspiration. The small loop represents a normal tidal volume. The peak expiratory flow (PEF) is labeled for the maximal expiratory curve. TLC = total lung capacity; RV = residual volume.
A family of flow-volume curves such as those depicted in Figure 2–23 is obtained in the same way as were the data in Figure 2–22, only in this case flow rates are plotted against lung volume for expiratory efforts of different intensities. Intrapleural pressures are not necessary. Because such curves can be plotted instantaneously, this test is often used clinically. There are 2 interesting points about this family of expiratory (upper portion) curves, which corresponds to the 3 curves in Figure 2–22. At high lung volumes, the airflow rate is effort-dependent, which can be seen in the left-hand portion of the curves. At low lung volumes, however, the expiratory efforts of different initial intensities all merge into the same effort-independent curve, as seen in the right-hand portion of the curve. Again, this difference is because intrapleural pressures high enough to cause dynamic compression are necessary to attain very low lung volumes, no matter what the initial expiratory effort. Also, at low lung volumes there is less alveolar elastic recoil pressure, and so there is less traction on the same airways and a smaller pressure gradient for airflow. Note that there is normally no effort independence on inspiratory curves if the subject breathes through the mouth. Patients with upper airway problems such as obstructive sleep apnea or vocal cord paralysis would demonstrate inspiratory effort independence.
The maximal flow-volume curve is often used as a diagnostic tool, as shown in Figure 2–24, because it helps distinguish between 2 major classes of pulmonary diseases—airway obstructive diseases and restrictive diseases, such as fibrosis. Obstructive diseases are those diseases that interfere with airflow; restrictive diseases are those diseases that restrict the expansion of the lung (see the pulmonary function test decision tree in Chapter 6).
Maximal expiratory flow-volume curves representative of obstructive and restrictive diseases.
Figure 2–24 shows that either obstruction or restriction can cause a decrease in the maximal flow rate that the patient can attain, the peak expiratory flow (PEF; shown in Figure 2–23), but that this decrease occurs for different reasons. Restrictive diseases, which usually entail elevated alveolar elastic recoil, may have decreased PEF because the TLC (and thus the VC) is decreased. The effort-independent part of the curve is similar to that obtained from a person with normal lungs. In fact, the FEV1/FVC is usually normal or even above normal since both the FEV1 and FVC are decreased because the lung has a low volume and because alveolar elastic recoil pressure may be increased. On the other hand, in patients with obstructive diseases, the PEF and FEV1/FVC are both low.
Obstructive diseases—such as asthma, bronchitis, and emphysema—are often associated with high lung volumes, which is helpful because the high volumes increase the alveolar elastic recoil pressure. The RV may be greatly increased if airway closure occurs at relatively high lung volumes. A second important feature of the flow-volume curve of a patient with obstructive disease is the effort-independent portion of the curve, which is depressed inward (concave): Flow rates are low for any relative volume.
Flow-volume curves are very useful in assessing obstructions of the upper airways and the trachea. Flow-volume loops can help distinguish between fixed obstructions (those not affected by the inspiratory or expiratory effort) and variable obstructions (changes in the transmural pressure gradient caused by the inspiratory or expiratory effort result in changes in the cross-sectional area of the obstruction). If the obstruction is variable, flow-volume loops can demonstrate whether the obstruction is extrathoracic or intrathoracic (Figure 2–25). A fixed obstruction affects both expiratory and inspiratory airflow (Figure 2–25A). Both the expiratory and inspiratory flow-volume curves are truncated, with decreased peak expiratory and peak inspiratory flows. The flow-volume loop is unable to distinguish between a fixed extrathoracic and a fixed intrathoracic obstruction, which would usually be determined with a bronchoscope. Fixed obstructions can be caused by foreign bodies or by scarring, usually from a previous intubation or tracheostomy, that makes a region of the airway too stiff to be affected by the transmural pressure gradient.
Inspiratory and expiratory flow-volume curves representing the patterns in: A: Fixed intra- or extrathoracic obstruction. B: Variable extrathoracic obstruction. C: Variable intrathoracic obstruction. Patm = atmospheric pressure; Paw = airway pressure; Ppl = intrapleural pressure; RV = residual volume; TLC = total lung capacity. (Reproduced with permission from Burrows B, Knudson RJ, Quan SF, Kettel LJ. Respiratory Disorders: A Pathophysiologic Approach. 2nd ed. Chicago: Year Book Medical Publishers; Copyright © 1983.)
During a forced expiration, the cross-sectional area of a variable extrathoracic obstruction increases as the pressure inside the airway increases (Figure 2–25B). The expiratory flow-volume curve is therefore nearly normal or not affected. However, during a forced inspiration, the pressure inside the upper airway decreases below atmospheric pressure, and unless the stability of the upper airway is maintained by reflex contraction of the pharyngeal muscles or by other structures, the cross-sectional area of the upper airway will decrease. Therefore, the inspiratory flow-volume curve is truncated in patients with variable extrathoracic obstructions. Variable extrathoracic obstructions can be caused by tumors, fat deposits, weakened or flabby pharyngeal muscles (as in obstructive sleep apnea), paralyzed vocal cords, enlarged lymph nodes, or inflammation.
During a forced expiration, positive intrapleural pressure decreases the transmural pressure gradient across a variable intrathoracic tracheal obstruction, decreasing its cross-sectional area and decreasing the PEF (Figure 2–25C). During a forced inspiration, as large negative intrapleural pressures are generated, the transmural pressure gradient across the variable intrathoracic obstruction increases and its cross-sectional area increases. Thus, the inspiratory flow-volume curve is nearly normal or not affected. Variable intrathoracic obstructions of the trachea are most commonly caused by tumors.
The dynamic compliance of the lungs is the change in the volume of the lungs divided by the change in the alveolar-distending pressure during the course of a breath. At low breathing frequencies, around 15 breaths/min and lower, dynamic compliance is about equal to static compliance, and the ratio of dynamic compliance to static compliance is 1 (Figure 2–26).
Illustration of changes in the ratio of dynamic compliance to static compliance with increasing breathing frequencies. The ratio changes little in normal subjects but decreases dramatically in patients with obstructive diseases of the small airways.
In normal persons, this ratio stays near 1 even at much higher breathing frequencies. However, in patients with elevated resistance to airflow in some of their small airways, the ratio of dynamic compliance to static compliance falls dramatically as breathing frequency is increased. This indicates that changes in dynamic compliance reflect changes in airways resistance as well as changes in the compliance of alveoli.
The effects of increased breathing frequency on dynamic compliance can be explained by thinking of a pair of hypothetical alveoli supplied by the same airway. Consider the time courses of their changes in volume in response to an abrupt increase in airway pressure (a “step” increase) in a situation in which the compliance of each alveolus or the resistance in the branch of the airway supplying it can be arbitrarily altered.
If the resistances and compliances of the 2 units were equal, the 2 alveoli would fill with identical time courses. If the resistances were equal, but the compliance of one were half that of the other, then it would be expected that the alveoli would fill with nearly identical flow rates but that the less compliant one would receive only half the volume received by the other. If the compliances of the 2 units were equal but one was supplied by an airway with twice the resistance to airflow of the one supplying the other, then it would be expected that the 2 units would ultimately fill to the same volume. However, the one supplied by the airway with elevated resistance would fill more slowly than the other because of its elevated resistance. (High resistance units take longer to fill because of lower flow rates; high compliance units take longer to fill because they can hold more volume.) This difference means that at high breathing frequencies the one that fills more quickly than the other will accommodate a larger volume of air per breath.
This situation may also lead to a redistribution of alveolar air after the inflating pressure has ceased because one alveolus has more air in it than the other. But both have equal compliance characteristics. The more distended one therefore has a higher elastic recoil pressure, and because they are joined by a common airway, some air is likely to follow the pressure gradient and move to the other.
Now let’s extrapolate this 2-unit situation to a lung with millions of airways supplying millions of alveoli. In a patient with small-airways disease, many alveoli may be supplied by airways with higher resistance to airflow than normal. These alveoli are sometimes referred to as “slow alveoli” or alveoli with long “time constants.” (The time constant is equal to the resistance times the compliance and it represents the time it takes for the alveolus to fill to 63% of its final volume.) As the patient increases the breathing frequency, the slowest alveoli will not have enough time to fill and will contribute nothing to the dynamic compliance. As the frequency increases, more and more slow alveoli will drop out and dynamic compliance will continue to fall. If this patient were being mechanically ventilated, alveoli may not have enough time to fill or empty (expiration is passive and mainly dependent on alveolar elastic recoil) between breaths. In the latter case “new” air may enter the lung before it has time to empty (“auto PEEP” or “stacked breaths”).