The movement of gases in and out of the respiratory system may be described by the physical laws that govern pressure, volume, and flow of gas. The study of these relationships is called “respiratory mechanics,” and the study of the energy cost of gas movement is called “energetics . ”
This chapter reviews basic terminology, respiratory mechanics under static and dynamic conditions, and energetics, including measurements of the work of breathing. The chapter also highlights how physiologic principles can be applied to specific clinical problems.
Flow is defined as the volume of gas passing a fixed point per unit of time. Flow is usually measured with a pneumotachygraph (flowmeter), which consists of a tube with a known fixed resistance. Flow can be calculated by measuring the pressure drop across the resistor.
Gas velocity is the distance moved by a gas molecule per unit of time and should not be confused with flow. At a constant flow, gas velocity will be greater in narrower tubes ( Fig. 5-1 ).
Volume is defined by the space occupied by a gas. The volume occupied by a fixed number of gas molecules is determined by temperature and pressure.
The volume of gas entering and leaving the lung can be determined by a spirometer that measures volume displacement or by integrating the flow signal measured by a pneumotachygraph. The subdivisions of lung volumes are shown in Figure 5-2 . Some of these subdivisions can be measured by spirometry alone (vital capacity, tidal volume), whereas others require use of plethysmography or helium dilution. Total lung capacity (TLC) is the lung volume at the end of a maximal inspiration. Residual volume is the volume at the end of a maximal expiratory effort. Functional residual capacity (FRC) refers to the volume in the lung at the end of a normal tidal exhalation, when there is normally relaxation of both inspiratory and expiratory muscles.
Total gas volume in the lung is usually measured at FRC using plethysmographic methods or by inert gas dilution methods, and each of these techniques has advantages and disadvantages. Traditional body plethysmography is widely available and accurate. For the body plethysmography technique, subjects are completely enclosed in a gas-tight container. Lung volume can be calculated by comparing changes in alveolar pressure (measured at the mouth while the patient pants against an occluded mouthpiece) with changes in pressure in the container. Body plethysmography may overestimate lung volumes because it will include measurement of abdominal gas if that gas is compressed and decompressed during the panting maneuver. Plethysmographs that use electrical inductance or optical data have been advanced as noninvasive methods for the measurement of lung volumes in those not able to tolerate traditional plethysmography. However, given the limited experience with these modalities in clinical settings, they are best considered experimental advances at this time.
Because spirometry cannot estimate FRC, one of several inert gas dilution techniques (usually with helium or nitrogen) or radiologic assessments is used. Historically, inert gas methods have usually used helium. For this method, subjects at FRC inhale a known concentration and volume of helium. The helium mixes with and is diluted by the gas already in the lung. A sample of exhaled gas is analyzed for helium concentration, allowing calculation of the FRC thus:
C 1 × V 1 = C 2 × ( V 1 + FRC )
FRC = V 1 × ( C 1 − C 2 ) /C 2
Another method uses the respiratory gases naturally present in a subject’s lungs (such as nitrogen) to be used for FRC determination. In the simplest form, measurement of FRC using the nitrogen dilution method requires the subject to be ventilated with a 100% oxygen gas mixture to allow the washout of nitrogen from the respiratory system. Once all the nitrogen has been washed out, the initial gas mixture (oxygen/air) is reinstituted, and the concentration of nitrogen is repeatedly sampled until it returns to the baseline level. The trajectory of nitrogen concentration in the exhaled gas allows the estimation of FRC. Unfortunately, the basic technique is complicated by time delays from gas analyzers. Adjustment for these delays is possible but is complex because the delay depends on gas viscosity, which changes as the nitrogen and oxygen concentrations change.
Imaging modalities such as computed tomography (CT) and magnetic resonance imaging can provide accurate assessment of FRC. CT may overestimate the volume of alveolar gas available for gas exchange because it measures the entire volume of gas in the lung irrespective of whether that gas is trapped and is not partaking in gas exchange.
The pressure of a gas is generated from the momentum of molecules colliding against a surface and is expressed as the force per unit area. Pressure, as opposed to force, is the same in all directions. Respiratory system pressures are usually reported relative to atmospheric pressure.
The pressures relevant to the respiratory system are shown in Figure 5-3 . Although it is possible to measure pleural pressure directly, it is usual to use esophageal pressure as a less invasive surrogate. Esophageal pressure may be measured using an air-filled balloon inserted into the middle third of the esophagus, approximately 35 to 45 cm from the nares. Balloons placed in the proximal third can be affected by movement of the neck, and balloons placed in the distal third can record falsely high readings owing to compression by the heart. Absolute values of pressure may be difficult to interpret in the supine position owing to the weight of the mediastinum.
The Combined Gas Law
For an ideal gas the relationship between volume, pressure, temperature, and the number of molecules of gas is described by the combined gas law:
PV = nRT
This equation summarizes earlier gas laws, including Boyle’s law (PV = constant), Charles’ law (V/T = constant), Gay-Lussac’s law (P/T = constant), and Avogadro’s law (n/V = constant). The volume of a gas should thus increase as the number of molecules of gas increases, the temperature increases, or the applied pressure decreases. Although this equation theoretically applies only to an ideal or perfect gas, it is an adequate approximation for clinical purposes.
Compliance, Resistance, and Time Constants
The respiratory system can be thought of as a combination of balloons and tubes ( Fig. 5-4 ). As the pressure across the wall of a balloon increases (called “transmural pressure,” or “pressure inside minus pressure outside”), the volume inside the balloon increases (see Fig. 5-4 ). The ratio between the change in volume and the change in transmural pressure is the compliance of the balloon. The units of compliance are volume/pressure. A large compliance indicates that volume changes markedly for every change in pressure. The inverse of compliance is elastance, with a greater value indicating a stiffer system.
Resistance is a measure of the pressure required to generate flow through a tube. The narrower the tube, the greater the pressure needed and the higher the resistance. The units of resistance are pressure/flow. The inverse of resistance is conductance, with a greater value indicating a more dilated airway. Flow in a tube may be laminar or turbulent, with significant implications for the amount of pressure drop or energy required to drive gas movement.
When a pressure is applied to a lung unit ( Fig. 5-5 ), the time required to fill the unit is dependent on its compliance and resistance. That is, it will take longer to fill if resistance is high because the flow will be reduced. Similarly, it will take longer to fill the unit if compliance is high because it will require more volume. The product of the resistance and compliance (RC) of a lung unit is the time constant and represents the time required for the lung unit to fill to 63% of the final volume if a constant pressure is applied. Areas of the lung with small time constants (i.e., low resistance and/or low compliance) will fill more rapidly than areas with larger time constants.
The effect of variations in time constants of lung units is shown in Figure 5-5 . This concept has clinical implications when there are heterogeneous time constants between parallel lung units, as in patients with chronic airflow obstruction or with patchy alveolar edema/atelectasis as is found in acute respiratory distress syndrome (ARDS). As long as respiratory frequency is slow, all units will fill to their static equilibrium. However, if respiratory rate increases so that the inspiratory time becomes less than the time constant of some units, these units will receive less ventilation and contribute to ventilation-perfusion mismatch.
Respiratory Mechanics in Static Conditions
We first consider the lung during static conditions, during which measurements are made while the respiratory system is maintained at a fixed volume with no gas flow. Even under static or no-flow conditions, pressure gradients are still required to distend the respiratory system (analogous to the balloon described previously). The energy used to distend the respiratory system during inhalation (elastic work) is stored as potential energy. Because of this stored energy, normal exhalation does not require work and is a passive process. This contrasts with resistive work (discussed later in this chapter), which cannot be stored as potential energy and is dissipated as heat. The elastic work performed to inflate the respiratory system has both lung and chest wall components.
Elastic Recoil of the Lungs
If removed from the body, an isolated lung will deflate because of its elastic recoil to a minimal volume containing only trapped gas. Because airways close before complete alveolar emptying, application of a negative pressure to the airway opening will not remove the trapped gas. The relationship between lung volume and transpulmonary pressure during inflation is shown in Figure 5-6 . Lung compliance is fairly high at the lung volumes associated with normal breathing but then decreases markedly near TLC. At TLC, very large increases in transmural pressure result in small changes in volume.
Two major factors are responsible for the elastic recoil of the lung: (1) lung connective tissue and (2) surface tension related to the air-liquid interface of the alveolar surface.
Lung Connective Tissue
A network of connective tissue fibers provides the framework for the alveoli and structural integrity for the lung. The fibers are composed mostly of collagen and elastin. Collagen fibers exhibit high tensile strength but are relatively noncompliant. They can be extended by only 2% of their length. This contrasts with elastin, which has a lower tensile strength and is more compliant; elastin fibers can be stretched by as much as 130% of their length. Experiments using selective destruction of fibers by collagenase or elastase show that elastin fibers are the major contributors to the volume-pressure relationships at low lung volumes, whereas collagen is more important at higher volumes approaching TLC. At low lung volumes, elastin fibers bear much of the stress, and collagen fibers are curled and unstressed. As lung volume increases, collagen fibers uncurl and straighten and have a major stiffening effect on the lung. In other words, the major role of collagen is to limit overdistention of the lung, and the major role of elastin is to facilitate inflation while providing lung stability to maintain the configuration of internal structures.
Destruction of lung connective tissue, from smoking-induced emphysema for instance, can substantially increase lung compliance. Contraction of smooth muscle in the airways and alveolar ducts may also affect elastic recoil because contraction of smooth muscle reduces volume by exerting traction on the lung (similar to a drawstring). Although contraction of airway smooth muscle in the peripheral lung units (small airways and alveolar ducts) changes lung compliance in several different animal species, there is little effect of contraction of lung smooth muscle on the shape of the volume-pressure curve in humans.
Alveolar Surface Forces and Surfactant
When a lung is completely filled with water, compliance is much greater than when it is filled with air. This suggests that the majority of lung elastic recoil is due to surface tension at the air-liquid interface lining the alveoli, rather than the recoil from elastin and collagen connective tissue fibers (see Chapter 8 for more details).
Because there are intermolecular forces of attraction between molecules of liquids, the surface of an air-liquid interface is under tension. That is, a molecule in the interior of a liquid is subjected to an equal force in all directions by other liquid molecules. However, the molecules at the surface are pulled toward each other and by molecules below the surface, and these forces (surface tension) cause a tendency for the surface to collapse.
Because of surface tension, a positive pressure must be present to prevent the collapse of a gas-filled bubble in a liquid. The pressure of gas (P) within a bubble is related to the surface tension (T) and the radius of curvature of the bubble (r) by Laplace’s law (P = 2T/r) ( Fig. 5-7 ). The surface of the alveolus with an air-liquid interface can be likened to a bubble with a radius of approximately 0.1 mm at FRC. If the surface tension were similar to that of water (72 mN/m), the lungs would be very noncompliant and ventilation with transpulmonary pressures in the physiologic range (3 to 5 cm H 2 O) would be impossible. Furthermore, the high surface tension would lead to alveolar instability and cause alveolar collapse. That is, smaller alveoli with smaller radii of curvature, by the nature of Laplace’s law, would have higher alveolar pressures than larger alveoli with greater radii of curvature. Because gas travels from an area of higher pressure to one of lower pressure, smaller alveoli would tend to empty into larger alveoli, eventually resulting in one large air-filled alveolus with all other alveolar units collapsed.
A number of factors contribute to alveolar stability in normal lungs. First, the connective tissue scaffolding limits overdistention of alveoli. Second, alveoli are interconnected (alveolar interdependence) by common walls and structures. Therefore, if one alveolus began to collapse, it would stretch adjacent alveolar walls, creating a tethering effect on the collapsing unit. Third, surfactant has an important stabilizing effect.
Hysteresis and Stress Adaptation
When normal lungs are inflated and deflated slowly, the volume-pressure curves are not identical. During inflation of the lung, the pressure required at any given lung volume is greater than that during deflation ( Fig. 5-8 ). The difference between the inflation and the deflation curves is due to hysteresis and stress adaptation.
Hysteresis refers to changes in mechanical properties due to the volume history of the lung and is caused by several factors. First, the effect of surfactant on surface tension is dependent on volume history, with surfactant less effective at reducing surface tension during inspiration than during expiration. During inspiration the phospholipids of surfactant move to the surface of the liquid layer, and during exhalation the surface film is compressed, the phospholipids are concentrated, and surface tension falls. Second, much higher pressures are required to open collapsed airways or alveoli than are required to keep them open. In disease states characterized by collapse of alveoli (e.g., ARDS), significant hysteresis of the pressure volume curve can be seen (see Fig. 5-8 ). This has clinical significance in patients with ARDS because higher inflation pressures are required to ventilate the lungs.
In contrast, stress adaptation refers to the time dependence of mechanics measurements. When lung tissue is stretched to a particular length, the tension required to maintain the length gradually diminishes due to time-dependent properties of surfactant and deformation of viscoelastic tissues of the lung. Therefore, after a sustained lung inflation, the pressure required to keep the lung at that volume will decrease.
Recruitment Maneuvers in Acute Respiratory Distress Syndrome
ARDS is a disease known to have surfactant dysfunction. The dysfunction of surfactant increases surface tension at the air-liquid interface of the alveoli, leading to collapse, intrapulmonary shunting, and hypoxemia. Because of alveolar collapse, hysteresis may be especially apparent in patients with ARDS (see Fig. 5-8 ). The lung collapse necessitates high inflation pressures to ventilate the lungs, thereby increasing the work of breathing and potentially leading to barotrauma; low tidal volumes are used in ARDS to mitigate further lung injury.
Due to hysteresis the pressure required to open an alveolus (“opening pressure”) is significantly greater than the pressure at which an already open alveolus will collapse (“closing pressure”). Because of this, “recruitment maneuvers” have been used in ventilated patients with ARDS who remain hypoxemic despite moderate levels of externally applied positive end-expiratory pressure (PEEP). A recruitment maneuver consists of a sustained inflation at a constant high pressure, which opens collapsed alveoli. Because closing pressures are much less than opening pressures, the alveoli stay open (at least temporarily) after the sustained inflation is ended, as long as the lung is not allowed to return to a very low volume ( Fig. 5-9 ).
Elastic Recoil of the Chest Wall
The chest wall is made up of the rib cage laterally, the sternum anteriorly, the vertebral column posteriorly, and the diaphragm caudally. Movement of the chest wall by the muscles of ventilation generates pressure gradients between the alveoli and the surrounding air, enabling gas to be moved in and out of the lungs.
Respiratory Muscles (see Chapter 97 )
The diaphragm consists of two separate muscles joined by a central tendon. The crural diaphragm arises from the first three lumbar vertebrae and the medial and lateral arcuate ligaments. The costal diaphragm arises from the inner surfaces and upper margins of the lower six ribs and sternum. Both costal and crural fibers insert onto the central tendon. The phrenic nerve innervates the diaphragm, with costal fibers innervated from the 3rd and 4th cervical spinal segments, and crural fibers from the 4th and 5th segments. In normal subjects the lower part of the rib cage encloses the upper part of the abdomen at FRC. In this area the diaphragm is pushed against the lower rib cage (zone of apposition) with fibers arranged in a cranial-caudal direction ( Fig. 5-10 ). Contraction of the diaphragm leads to a piston-like up-and-down motion of the diaphragm. Furthermore, diaphragm contraction leads to an increase in abdominal pressure, which pushes the rib cage upward and outward.
In patients with substantial hyperinflation (e.g., emphysema), the position of the diaphragm is lower, with loss of the zone of apposition. When the zone of apposition is lost, contraction of the diaphragm is less mechanically advantageous and can result in paradoxical inward movement of the lower rib cage during inhalation (the Hoover sign) and less piston-like motion of the diaphragm dome. Furthermore, the length of the diaphragm is shortened, which creates a further mechanical disadvantage because the ability of the muscle to generate force is attenuated at shorter lengths ( Fig. 5-11 ). However, in animals and humans, compensatory shortening of the diaphragm through loss of sarcomeres has been shown to mitigate these effects and help to improve the length-tension characteristics of the diaphragm. The improved symptoms experienced by patients with chronic obstructive pulmonary disease (COPD) following lung volume reduction surgery or after removal of pleural effusions may be related to increased length of the zone of apposition and improved mechanical function of the diaphragm.
During inhalation the rib cage moves laterally and anteriorly. The anterior motion of the sternum is akin to a “pump-handle,” motion, and the lateral movement and elevation of the ribs is akin to a “bucket-handle” motion ( Fig. 5-12 ). The muscles of the rib cage include the internal intercostal muscles, the external intercostal muscles, and the parasternal muscles. Internal intercostal muscles are used mainly for exhalation, whereas external muscles are used for inhalation. The parasternal muscles are located next to the sternum; although they are part of the internal intercostal muscle layer, their major function is to raise the sternum during inhalation.
Other muscles used in inhalation include the scalene muscles, which help to elevate the rib cage and prevent paradoxical inward movement of the upper rib cage with diaphragm displacement. Under conditions of ventilatory stress, other muscles not normally used for respiration can be recruited. These include the sternocleidomastoid, pectoralis, trapezius, and muscles of the vertebral column. Contraction of the muscles in the abdominal wall can be used to aid exhalation.
Fatigue of Respiratory Muscles
Fatigue develops when the rate of work required by the muscles exceeds the rate of energy supply. Usually there is substantial reserve of the respiratory muscles that prevents fatigue. However, under conditions of excessive elastic or resistive load (either experimentally induced or due to disease states) and/or inadequate energy supply (e.g., during hypoxemia or cardiogenic shock), fatigue develops and impairs respiratory muscle performance. Recovery from fatigue usually requires more than 24 hours of rest.
One measure of workload of the diaphragm is the pressure-time index. This is the product of the percentage of maximal pressure (i.e., transdiaphragmatic pressure/maximal transdiaphragmatic pressure) exerted by the muscle multiplied by the percentage of time spent during inspiration (inspiratory time/total time). In normal subjects a value greater than 15% to 18%, if sustained, will lead to fatigue and loss of force-generating capacity of the diaphragm.
Central respiratory fatigue arises when there is a reduction in central motor output to the muscles. Central fatigue can be identified when direct electrical stimulation of respiratory muscles produces more muscle force generation than the patient generates with a maximal voluntary effort. This decrease in central motor output may represent an attempt to prevent damage to the muscles under conditions of inadequate energy supply. Peripheral fatigue represents failure at or past the level of the neuromuscular junction. With peripheral fatigue, force generation is reduced despite a constant electrical stimulation. Peripheral fatigue may arise because of failure of transmission at the neuromuscular junction, reduced adenosine triphosphate level, reduced calcium availability from the sarcoplasmic reticulum, or reduced calcium sensitivity of the myofilaments (caused by acidosis or increased inorganic phosphate).
Respiratory Muscle Atrophy during Positive-Pressure Mechanical Ventilation
Although excessive loading of the respiratory muscles can lead to fatigue and a reduction in force generation, excessive unloading of the muscle may lead to muscle atrophy. Mechanical ventilation in paralyzed rats and baboons leads to diaphragmatic atrophy. Furthermore, human studies in organ donors have shown significant diaphragmatic atrophy after 18 to 69 hours of diaphragmatic inactivity. However, the clinical significance of disuse atrophy in ventilated patients is still unclear. Patients who are not paralyzed may still have substantial activation of respiratory muscles while receiving mechanical ventilation, and this degree of stimulation may be adequate to prevent atrophy. Furthermore, there are other potential causes of muscle weakness or injury in the intensive care unit, including electrolyte disturbances, malnutrition, medications and toxins (e.g., alcohol, corticosteroids, aminoglycosides), sepsis-related cytokines, and inflammatory disorders (e.g., polymyositis), so that attributing muscle weakness to disuse atrophy of muscle fibers in the clinical setting is difficult.
Chest Wall Compliance (see Chapter 98 )
The compliance of the chest wall can be determined by measuring changes in the transmural pressure across the thorax (pleural pressure minus body surface pressure) relative to change in volume. The volume-pressure curve of the relaxed chest wall is shown in Figure 5-6 . In contrast to the lung, the resting volume of the relaxed chest wall is approximately 1 L above FRC. The difference between the resting volume of the lung and chest wall can be appreciated in patients with a pneumothorax, in which the chest wall recoils outward and the lung recoils inward. When the chest wall is distended above this resting volume, the relaxed chest wall recoils inward, and when the chest wall is moved below this resting volume, it recoils outward. The compliance of the normal chest wall is approximately 200 mL/cm H 2 O, but it becomes progressively stiffer at lower lung volumes, and it is this stiffness that predominantly determines residual volume in normal young subjects.
Factors such as ossification of costal cartilage, arthritis of costovertebral joints, skin eschars from burns, obesity, and abdominal distention may reduce chest wall compliance. Chest wall compliance increases in the sitting position compared with the supine position, but the overall effect is usually modest.
Integration of Lung and Chest Wall Mechanics
The chest wall and lung are juxtaposed and, in the absence of pleural disease (such as pneumothorax or pleural fibrosis), changes in chest wall volume are essentially identical to changes in lung volume ( Fig. 5-13 ). An integration of lung and chest wall mechanics can be graphically represented by plotting the pleural pressure against the relaxed volumes of the lung and chest wall (Campbell diagram) ( Fig. 5-14 ). These plots help to explain the determinants of the commonly measured static lung volumes. At FRC (relaxation volume), the outward recoil pressure of the chest wall is balanced by the inward recoil of the lung. The pleural pressure at FRC is negative (approximately −3 to −4 cm H 2 O) in normal subjects. Changes in the compliance of the lung or chest wall may change the relaxation volume (see Fig. 5-14 ).
Increases in lung compliance (e.g., due to emphysema) will tend to shift the volume-pressure curve of the lung upward and to the left. If chest wall compliance remains constant, FRC will increase. In contrast, if lung compliance is reduced (e.g., due to pulmonary fibrosis or ARDS), the volume-pressure curve of the lung will shift downward and to the right, leading to a reduction in FRC. Similarly, changes in chest wall compliance can lead to increases or decreases in FRC.
FRC is believed to represent the balance of forces in a relaxed state at which there is no muscle activation. However, FRC may not be completely passively determined. There is a reduction in FRC with paralysis compared with the relaxed state, suggesting that inspiratory muscle tone of the rib cage muscles and diaphragm may contribute to “passive” chest wall compliance and thus to FRC.
With activation of inspiratory and expiratory muscles, the configuration of the chest wall changes. This is demonstrated in Figure 5-15 . Activation of inspiratory muscles effectively shifts the pressure-volume curve of the chest wall, so that at any given pleural pressure, the chest wall volume is increased relative to the relaxed state. The horizontal difference between the relaxed curve and the curve with inspiratory muscle activation represents the net pressure generated by the inspiratory muscles.
TLC is determined by the balance between the strength of the inspiratory muscles counterbalanced by the inward lung recoil plus the inward chest wall recoil. However, in normal subjects, TLC is predominantly determined by the markedly increased stiffness of the lung at high lung volume, rather than the recoil of the chest wall or respiratory muscle strength. In support of this concept, inspiratory muscle training in normal subjects increases strength substantially (55% increase in maximal inspiratory pressure) but results in a very modest change in TLC and VC (about 4%). The situation in patients with neuromuscular disease is very different; in such individuals, reduced TLC is largely due to respiratory muscle weakness and can be improved to some extent by inspiratory muscle training.
Activation of expiratory muscles shifts the pressure-volume curve of the chest wall in the other direction, and the degree of shift is a measure of the net pressure exerted by the expiratory muscles. In young subjects, residual volume is determined by the balance of the strength of the expiratory muscles counterbalanced by the outward recoil of the chest wall and inward lung recoil (small component).
In older subjects or patients with obstructive lung disease, residual volume is not determined by this balance of forces but is rather determined by airway closure. In such individuals, airways close and gas is trapped during exhalation before a static balance is achieved.
Calculation of Total Respiratory System Compliance from Lung and Chest Wall Compliance
The chest wall and lung act as capacitors in series, analogous to electrical capacitors. The relationship between total respiratory system compliance and lung and chest wall compliance is indicated by the following formula:
1 /Total compliance = 1 /Lung compliance + 1 /Chest wall compliance
Alternatively, since elastance = 1/compliance,
Total elastance = Elastance of lung + Elastance of chest wall
As an example, in a supine paralyzed subject, lung compliance is approximately 150 mL/cm H 2 O pressure, and chest wall compliance is approximately 200 mL/cm H 2 O pressure ; the calculated compliance of the respiratory system would thus be approximately:
1 / ( 1 / 200 + 1 / 150 ) = 85.7 mL/cm H 2 O
The Impact of Positive End-Expiratory Pressure on Pleural Pressure
In patients receiving positive-pressure mechanical ventilation for diseases that decrease lung compliance, pressure is usually applied at the end of exhalation to prevent alveolar collapse. Some—but not all—of this PEEP is transmitted from the alveolar space to the pleural space (see Fig. 5-13 ). This is important because pericardial pressure will increase to a similar extent as pleural pressure. An increase in pericardial pressure reduces venous return and can lead to reduced cardiac output and hypotension. Furthermore, PEEP will increase central venous pressure and pulmonary artery occlusion pressure to a magnitude similar to the increase in pleural pressure. If this is not taken into account, errors may be made when using these pressures to assess patients’ intravascular volume status.
During mechanical ventilation, changes in both pleural pressure and transpulmonary pressure depend on the relationships between the elastance of the chest wall (Ew), the lung (E l ), and the total respiratory system (Ers, which is the sum of Ew and E l ), a relationship (see Fig. 5-13 ) that may be described by the equation:
where Ppl is the pleural pressure and P aw is the flow-resistive pressure in the airway.
In healthy individuals, Ew and E l are approximately equal, and Ew/Ers has a value of approximately 0.5 and remains constant within the range of normal ventilation. In this situation it is reasonable to infer the value of Ppl from P aw . However, in disease, both Ew and E l demonstrate great variability, such that the ratio of Ew to Ers is unpredictable. In this situation, Ppl cannot be assumed to be directly related to P aw . If lung compliance is low and chest wall compliance is high, very little of the alveolar pressure will be transmitted to the pleural space and vice versa. In healthy individuals the chest wall and lung compliances are equal, and thus the increase in pleural pressure is roughly equal to half of the applied PEEP (e.g., 5 cm H 2 O pressure if the PEEP is 10 cm H 2 O pressure).
However, if lung compliance is much less than chest wall compliance (e.g., in a thin patient with severe ARDS), the pleural pressure will change minimally with PEEP. In contrast, if lung compliance is much greater than chest wall compliance (e.g., in a patient with emphysema and kyphoscoliosis), a much greater proportion of PEEP will be transmitted to the pleural space.
Plateau Pressures in Patients Receiving Positive-Pressure Mechanical Ventilation
In patients receiving positive-pressure mechanical ventilation, the plateau pressure is the distending pressure of the respiratory system at end inspiration, when there is no flow. When possible, plateau pressures should be limited to a maximum of 30 to 35 cm H 2 O because higher pressures can damage the lung through overdistention. This recommendation is based in part on animal studies that show that pressures below this range seem to protect the lungs from injury. There is also a physiologic basis for this practice because at TLC in a normal subject, the distending pressure across the lung (alveolar pressure minus pleural pressure) is approximately 35 cm H 2 O. Therefore limiting the plateau pressure to less than 35 cm H 2 O should limit lung expansion to below TLC and not overdistend the lung. The use of low tidal volumes during mechanical ventilation for lung injury has led to improved clinical outcomes, presumably because of decreased alveolar distention and volutrauma. In a landmark study, tidal volumes were limited to 4 to 6 mL/kg of ideal body weight in an attempt to limit lung overdistention. However, others have argued that alveoli may become overdistended and undergo cyclic recruitment-derecruitment even with low tidal volumes in some patients, whereas ultralow tidal volumes (4 mL/kg of ideal body weight or less) may decrease this phenomenon further.
Although limiting plateau pressures and tidal volumes during mechanical ventilation has demonstrated some benefit in decreasing ventilator-associated lung injury, there is no clear limit below which further decreases will not improve outcomes. This may be because neither plateau pressure nor tidal volume measurements account for the effect of pleural pressure on lung mechanics. During mechanical ventilation, plateau pressure is often used as a surrogate for transpulmonary pressure (pressure used to distend the lung), when it is actually a measure of transrespiratory system pressure. When lung compliance is much less than chest wall compliance, this is a reasonable assumption, because during mechanical ventilation, pleural pressure will not be substantially greater than body surface pressure. However, in conditions in which the chest wall is stiff (e.g., obesity, after abdominal surgery), the plateau pressure may substantially overestimate the distending pressure of the lung, because pleural pressure may be much greater than body surface pressure. Indeed, some investigators have suggested that measurement of pleural pressure (with an esophageal balloon) could be useful in titrating mechanical ventilation in ARDS.
Stress and Strain
To an engineer, stress is the net difference between forces acting on a body (in this case, the lung parenchyma). From the point of view of lung mechanics, stress reflects the difference between distending forces (for example, alveolar pressure) and collapsing forces (the pleural pressure in this case), and stress is therefore represented by the transpulmonary pressure.
The deformation of a structure by stress is called “strain,” which is defined as the change in size or shape compared to the structure’s initial shape or volume. That is:
Strain = Δ V/FRC
For example, if a lung is inflated with a 500-mL tidal volume from an initial FRC of 1500 mL, the strain associated with this is 0.33 (500 mL/1500 mL). The recommendation of a specific tidal volume for lung-protective ventilation is predicated on the assumption that there is a uniform amount of lung able to receive the inspired volume. This volume is the amount of aerated lung at end expiration—the FRC. Underlying this is the assumption that the FRC can be predicted by ideal body weight—an assumption that is not correct in injured lungs. It is therefore understandable that, given the same ideal body weights and tidal volume prescriptions, two patients with ARDS might have significantly different lung strain if they have different FRCs. It therefore may be prudent to attempt to minimize strain during mechanical ventilation by measuring FRC and titrating tidal volume to this assessment rather than ideal body weight. Indeed, there is some evidence that maintaining strain less than 1.5 to 2 reduces markers of lung parenchymal injury. From a practical point of view, however, it is time consuming and laborious to measure FRC in a serial fashion. Although there are commercially available FRC measurement systems, either as stand-alone systems or integrated into mechanical ventilators, the cost of these systems is not insignificant. Fortunately, it is possible to estimate strain from measurements of stress because the two values are linked by a constant (the specific elastance of lung [E l spec ]) such that:
Strain = Stress/EL spec
Fortunately, the value of E l spec is fairly constant throughout health and disease, with a value of approximately 13 cm H 2 0. Thus by measuring stress, strain may be inferred (i.e., approximately one thirteenth of stress). Unfortunately, neither of these measures accounts for the local heterogeneity of mechanical characteristics found in injured lung. It is therefore possible that susceptible lung units are subject to critically injurious forces even when global measures of stress and strain appear acceptable.
The Respiratory System in Dynamic Conditions
The work to distend the chest wall and lung during inhalation is stored as potential energy (elastic work). Work also must be done to overcome the inertia of the gas (which is negligible when breathing air at normal breathing frequency and atmospheric pressures) and nonelastic resistance (resistive work), which cannot be stored, but is dissipated as heat. Resistive work has two components that must be considered.
Resistive Work Due to Gas Flow Through Airways
When gas flows through an airway, frictional and viscous forces cause energy loss and a pressure drop along the airway. The extent of this pressure drop is dependent on the resistance to gas flow, physical properties of the gas (e.g., density and viscosity), and the nature of flow (laminar versus turbulent).
Laminar versus Turbulent Flow
The frictional drop is least with pure laminar flow, in which gas molecules travel in a straight line. The velocity profile of the gas is parabolic, with the molecules closer to the wall traveling slower than molecules in the center of the airway ( Fig. 5-16 ). When gas flow is laminar, the flow is proportional to the pressure gradient (ΔP) along the airway and inversely proportional to resistance (R).
Flow = Δ P / R