Elastic Opposition to Ventilation

Elastin and collagen fibers are found in the lung parenchyma. These tissues give the lung the property of elasticity. Elasticity is the physical tendency of an object to return to an initial state after deformation. When stretched, an elastic body tends to return to its original shape. The tension developed when an elastic structure is stretched is proportional to the degree of deformation produced (Hooke’s law). An example is a simple spring (Figure 10-3). When tension on a spring is increased, the spring lengthens. However, the ability of the spring to stretch is limited. When the point of maximal stretch is reached, further tension produces little or no increase in length. Additional tension may break the spring.

In the respiratory system, inflation stretches tissue. The elastic properties of the lungs and chest wall oppose inflation. To increase lung volume, pressure must be applied. This property may be shown by subjecting an excised lung to changes in transpulmonary pressure and measuring the associated changes in volume (Figure 10-4). To simulate the pressures during breathing, the lung is placed in an airtight jar. The force to inflate the lung is provided by a pump that varies the pressure around the lung inside the jar, simulating P_pl. This action mimics the pleural pressure changes associated with thoracic expansion and contraction. The changes in transpulmonary pressure are made in discrete steps, allowing the lungs to come to rest in between so that all of the applied pressure opposes elastic forces and none of it opposes resistive forces (i.e., flow is zero when the measurements are made). The amount of stretch (inflation) is measured as volume by a spirometer. Changes in volume resulting from changes in transpulmonary pressure are plotted on a graph.

During inspiration in this model, increasingly greater negative pleural pressures are required to stretch the lung to a larger volume. As the lung is stretched to its maximum (total lung capacity [TLC]), the inflation “curve” becomes flat. This flattening indicates increasing opposition to expansion (i.e., for the same change in transpulmonary pressure, there is less change in volume).²

As with a spring when tension is removed, deflation occurs passively as pressure in the jar is allowed to return toward atmospheric. Deflation of the lung does not follow the inflation curve exactly. During deflation, lung volume at any given pressure is slightly greater than it is during inflation. This difference between the inflation and deflation curves is called hysteresis.² Hysteresis indicates that factors other than simple elastic tissue forces are present. The major factor, particularly in sick lungs, is the opening of collapsed alveoli during inspiration that tend to stay open during expiration until very low lung volumes are reached.

Chest wall and lung elastances are connected in series, meaning that they both experience the same flow and change in volume, but they do not have the same pressure differences. Series elastances are simply additive. The elastance of the respiratory system is the sum of lung (pulmonary) elastance (E_L) and chest wall elastance (E_CW):

ERS=EL+ECW

Expressed in terms of compliance:

CRS=CL×CCWCL+CCW

The resistance of the natural and artificial airways (e.g., endotracheal tube) is also in series so that the total system resistance is simply the sum of resistance of the components.

Surface Tension Forces

Part of the hysteresis exhibited by the lung is a result of surface tension forces in the alveoli. If a lung is filled with fluid such as saline, the pressure-volume curves look much different than the pressure-volume curves of an air-filled lung (Figure 10-5). Less pressure is needed to inflate a fluid-filled lung to a given volume. This phenomenon indicates that a gas-fluid interface in the air-filled lung changes its inflation-deflation characteristics.

The recoil of the lung is a combination of tissue elasticity and these surface tension forces in the alveoli. During inflation, additional pressure is needed to overcome surface tension forces. During deflation, surface tension forces are reduced, resulting in altered pressure-volume characteristics (i.e., the leftward shift seen in Figure 10-4). In the intact lung (i.e., within the chest), the volume history also affects the degree of hysteresis that occurs. Factors such as the initial volume, the tidal excursion, and whether the lungs have been previously inflated or deflated help determine the volume history and the shape of the pressure-volume curves of the lung.

A phospholipid called pulmonary surfactant reduces surface tension in the lung. Alveolar type II cells probably produce pulmonary surfactant (see Chapter 8). In contrast to typical surface-active agents, pulmonary surfactant changes surface tension according to its area.³ The ability of pulmonary surfactant to reduce surface tension decreases as surface area (i.e., lung volume) increases. Conversely, when surface area decreases, the ability of pulmonary surfactant to reduce surface tension increases. This property of changing surface tension to match lung volume helps stabilize the alveoli. Any disorder that alters or destroys pulmonary surfactant can cause significant changes in the work of distending the lung.

Lung Compliance

Tissue elastic forces and surface tension oppose lung inflation. Compliance is the reciprocal of elastance:

Compliance=1Elastance=ΔVΔP

Compliance is defined as volume change per unit of change in the pressure difference across the structure. It is usually measured in milliliters per centimeter of water.

A graph of change in lung volume versus change in transpulmonary pressure (Figure 10-6, A) is the compliance curve of the lungs. Figure 10-6, B compares a normal lung compliance curve with curves that might be observed in patients who have emphysema (obstructive lung disease) or pulmonary fibrosis (restrictive lung disease). The curve from a patient with emphysema is steeper and displaced to the left. The shape and position of this curve represent large changes in volume for small pressure changes (increased compliance). Increased compliance results primarily from loss of elastic fibers, which occurs in emphysema. The lungs become more distensible so that the normal transpulmonary pressure results in a larger lung volume. The term hyperinflation is used to describe an abnormally increased lung volume. A distinctly opposite pattern is seen in pulmonary fibrosis. Interstitial fibrosis is characterized by an increase in connective tissue. The compliance curve of a patient with pulmonary fibrosis is flatter than the normal curve, shifted down and to the right. As a result, there is a smaller volume change for any given pressure change (decreased compliance). Consequently, the lungs become stiffer, usually with a reduced volume.

Chest Wall Compliance

Inflation and deflation of the lung occur with changes in the dimensions of the chest wall (see Chapter 8). The relationship between the lungs and the chest wall can be illustrated by plotting their relaxation pressure curves separately and combined (Figure 10-7). In the intact thorax, the lungs and chest wall recoil against each other. The point at which these opposing forces balance determines the resting volume of the lungs, or functional residual capacity. This is also the point at which alveolar pressure equals atmospheric pressure. The normal FRC is approximately 40% of the TLC. If the normal lung–chest wall relationship is disrupted, the lung tends to collapse to a volume less than the FRC, and the thorax expands to a volume larger than the FRC.

The lung–chest wall system may be compared with two springs that are pulling against each other. The chest wall spring tends to expand, whereas the lung spring tends to contract. At the resting level, the forces of the chest wall and lungs balance. The tendency of the chest wall to expand is offset by the contractile force of the lungs. This balance of forces determines the resting lung volume, or FRC. The opposing forces between the chest wall and lungs are partially responsible for the subatmospheric pressure in the intrapleural space. Diseases that alter the compliance of either the chest wall or the lung often disrupt the balance point, usually with a change in lung volume.

Inhalation occurs when the balance between the lungs and chest wall shifts. Energy from the respiratory muscles (primarily the diaphragm) overcomes the contractile force of the lungs. At the beginning of the breath, the tendency of the chest wall to expand facilitates lung expansion. When lung volume nears 70% of the vital capacity, the chest wall reaches its natural resting level. To inspire to a lung volume greater than about 70% of TLC, the inspiratory muscles must overcome the recoil of both the lungs and the chest wall (see Figure 10-7).

For exhalation, potential energy “stored” in the stretched lung (and chest wall at high volumes) during the preceding inspiration causes passive deflation. To exhale below the resting level (FRC), muscular effort is required to overcome the tendency of the chest wall to expand. The expiratory muscles (see Chapter 8) provide this energy.

Compliance of the chest wall, similar to lung compliance, is a measure of distensibility. The compliance of the normal chest wall is similar to that of the lungs (0.2 L/cm H₂O). Obesity, kyphoscoliosis, ankylosing spondylitis, and many other abnormalities can reduce chest wall compliance and lung volumes.

Frictional (Nonelastic) Opposition to Ventilation

Frictional forces also oppose ventilation. Frictional opposition forces differ from the elastic properties of the lungs and thorax. Frictional opposition occurs only when the system is in motion. Frictional opposition to ventilation has the following two components: tissue viscous resistance and airway resistance.

Tissue Viscous Resistance

Tissue viscous resistance is the impedance of motion caused by displacement of tissues during ventilation. Displaced tissues include the lungs, rib cage, diaphragm, and abdominal organs. The energy to displace these structures is comparable to the impedance caused by friction in any dynamic system. Tissue resistance accounts for only approximately 20% of the total resistance to lung inflation. Obesity, fibrosis, and ascites can alter tissue viscous resistance, increasing the total impedance to ventilation.

Airway Resistance

Gas flow through the airways also causes frictional resistance. Impedance to ventilation by the movement of gas through the airways is called airway resistance. Airway resistance accounts for approximately 80% of the frictional resistance to ventilation.

Airway resistance is the ratio of driving pressure responsible for gas movement to the flow of the gas, calculated as follows:

Raw=ΔPTAΔV˙=PAO−PAΔV˙

where R_aw is resistance, P_TA is transairway pressure difference, is flow, P_AO is pressure at the airway opening, and P_A is alveolar pressure.

Driving pressure is measured in centimeters of water (cm H₂O), and flow is measured in liters per second (L/sec). Airway resistance (R_aw) is recorded in cm H₂O/L/sec or, more accurately, cm H₂O • sec • L⁻¹. Airway resistance in healthy adults ranges from approximately 0.5 to 2.5 cm H₂O/L/sec. To cause gas to flow into or out of the lungs at 1 L/sec, a healthy subject needs to lower his or her alveolar pressure only 0.5 to 2.5 cm H₂O below atmospheric pressure.

R_aw in nonventilated patients is usually measured in a pulmonary function laboratory (see Chapter 19). Flow () is measured with a pneumotachometer. Alveolar pressures are determined in a body plethysmograph, an airtight box in which the patient sits. By momentarily occluding the patient’s airway and measuring the pressure at the mouth, alveolar pressure can be estimated (i.e., mouth pressure equals alveolar pressure under conditions of no flow). By relating flow and alveolar pressure to changes in plethysmograph pressure, airway resistance can be calculated.

Factors Affecting Airway Resistance

Two patterns characterize the flow of gas through the respiratory tract: laminar flow and turbulent flow. A third pattern, tracheobronchial flow, is a combination of laminar and turbulent flow. When flow is laminar, gas moves in discrete layers, or streamlines. Layers near the center of an airway move faster than layers close to the wall of the airway; this results from friction between gas molecules and the wall.

Poiseuille’s law (see Chapter 6) defines laminar flow through a smooth, unbranched tube of fixed dimensions (i.e., length and radius). The pressure required to cause a specific flow of gas through a tube is calculated as follows:

ΔP=η8lV˙πr4

where:

ΔP = Driving pressure (dynes/cm²)

η = Coefficient of viscosity of the gas

l = Tube length (cm)

 = Gas flow (ml/sec)

r = Tube radius (cm)

(π and 8 are constants.)

By eliminating factors that remain constant, such as viscosity, length, and known constants, this equation can be rearranged as follows to solve for ΔP:

ΔP=V˙r4

This equation says that for gas flow to remain constant, delivery pressure must vary inversely with the fourth power of the airway’s radius. Reducing the radius of a tube by half requires a 16-fold pressure increase to maintain a constant flow. To maintain ventilation in the presence of narrowing airways, large increases in driving pressure may be needed. The energy necessary to generate these pressures can markedly increase the work of breathing.

 Rule of Thumb

A change in the caliber of an airway by a factor of 2 causes a 16-fold change in resistance. This rule applies to human airways and artificial airways (i.e., endotracheal and tracheostomy tubes). If the size of a patient’s airway is reduced from 2 mm to 1 mm, airway resistance increases by a factor of 16. Similarly, if a 4.5-mm endotracheal tube is replaced with a 9-mm tube, the pressure required to cause a flow of 1 L/sec through the tube decreases 16-fold. This rule has many practical consequences. It is the basis for bronchodilator therapy and for using the largest practical size of artificial airway.

Another way to express the relationship between pressure and flow is as follows:

V˙=ΔPr4

This equation shows that if the gas delivery pressure ventilating the lung remains constant, gas flow varies directly with the fourth power of the airway’s radius. Reducing the airway radius by half decreases the flow 16-fold at a constant pressure. Small changes in bronchial caliber can markedly change gas flow through an airway.

Under certain conditions, gas flow through a tube changes significantly. The orderly pattern of concentric layers is no longer maintained. Gas molecules form irregular currents. This pattern is called turbulent flow. Transition from laminar to turbulent flow depends on the following factors: gas density (d), viscosity (h), linear velocity (v), and tube radius (I). Table 10-3 compares changes in flow and pressure resulting from laminar and turbulent flow.

TABLE 10-3

Comparison of Driving Pressures, Laminar versus Turbulent Flow

Flow	Laminar	Turbulent
1	1	1
2	2	4
4	4	16
8	8	64
16	16	256

Values are nondimensional units showing proportional effect.

 Rule of Thumb

Patients who have emphysema can directly influence the EPP in their airways to reduce airway collapse and closure. Airway collapse may occur in patients who have emphysema because the normal support structure for small airways has been destroyed. By exhaling through “pursed lips,” a patient with emphysema changes the pressure at the airway opening. The gentle back pressure created counters the tendency for small airways to collapse by moving the EPP toward larger airways.

Distribution of Airway Resistance

Approximately 80% of the resistance to gas flow occurs in the nose, mouth, and large airways, where flow is mainly turbulent. Only about 20% of the total resistance to flow is attributable to airways smaller than 2 mm in diameter, where flow is mainly laminar. This fact seems to contradict the fact that resistance is inversely related to the radius of the conducting tube.

Branching of the tracheobronchial tree increases the cross-sectional area with each airway generation (Figure 10-8). As gas moves from the mouth to the alveoli, the combined cross-sectional area of the airways increases exponentially (see Chapter 8). According to the laws of fluid dynamics, this increase in cross-sectional area causes a decrease in gas velocity. The decrease in gas velocity promotes a laminar flow pattern, particularly in smaller (i.e., <2 mm) airways.

Turbulent flow predominates in the mouth, trachea, and primary bronchi (Table 10-4). Gas velocity is high in the bigger airways, favoring turbulent flow patterns. At the level of the terminal bronchioles, the cross-sectional area increases more than 30-fold. Gas velocity is very low here. In normal small airways, flow is laminar. The driving pressure across these airways is less than 1% of the total driving pressure for the system.

TABLE 10-4

Distribution of Airway Resistance

Location	Total Resistance (%)
Nose, mouth, upper airway	50
Trachea and bronchi	30
Small airways (<2 mm)	20

The diameter of the airways is not constant. During inspiration, the stretch of surrounding lung tissue and widening transpulmonary pressure gradient increase the diameter of the airways. The higher the lung volume, the more that these factors influence airway caliber (Figure 10-9). The increase in airway diameter with increasing lung volume decreases airway resistance. As lung volume decreases toward residual volume, airway diameters also decrease; this explains why wheezing (see Chapter 15) is most often heard during exhalation. Airway resistance increases dramatically at low lung volumes.

Differences in Thoracic Expansion

The conical configuration of the thorax and the action of the respiratory muscles cause proportionately greater expansion at the lung bases than at the apices (see Chapter 8). Expansion of the lower chest is approximately 50% greater than expansion of the upper chest.²¹ The action of the normal diaphragm preferentially inflates the lower lobes of the lung.

Transpulmonary Pressure Gradients

The transpulmonary pressure gradient (see earlier section on Mechanics of Ventilation) is not uniform throughout the thorax. It varies substantially within the lung and from the top to the bottom of the lung. At a given level of alveolar inflation, the transpulmonary pressure gradient is directly related to the pleural pressure. Pleural pressure represents the pressure on the outer surface of the lung. Its effect lessens toward more centrally located alveoli. Changes in the transpulmonary pressure gradient are greatest in peripheral alveoli (i.e., near the surface of the lung). The changes are least in the alveoli of the central zones. Peripheral alveoli expand proportionately more than their more central counterparts.

Top-to-bottom differences in pleural pressure have an even greater effect on the distribution of ventilation, especially in the upright lung.³ Pleural pressure increases by approximately 0.25 cm H₂O for each 1 cm, from the lung apex to its base for the average-sized adult lung. This increase in pressure results from the weight of the lung itself and the effect of gravity. In an adult-sized lung (approximately 30 cm from apex to base), pleural pressure at the apex is approximately −10 cm H₂O. At the base, pleural pressure is only about −2.5 cm H₂O. Because of these differences, the transpulmonary pressure gradient at the top of the upright lung is greater than it is at the bottom. As a result, alveoli at the apices have a larger resting volume than do alveoli at the bases.

Because of their larger volume, alveoli at the apices expand less during inspiration than alveoli at the bases. Apical alveoli rest on the upper portion of the lung’s pressure-volume curve (Figure 10-18). This part of the curve is relatively flat. Each unit of pressure change causes only a small change in volume. Alveoli at the lung bases are positioned on the steeper middle portion of the pressure-volume curve. For each unit of pressure change, there is a larger change in volume (greater compliance). For a given transpulmonary pressure gradient, alveoli at the bases expand more than alveoli at the apices. The bases of the upright lung receive approximately four times as much ventilation as the apices.

These gravity-dependent differences also are observed in recumbent individuals. The magnitude of the differences is less than in the upright lung because the top-to-bottom distance is less. Ventilation is still greatest in the dependent zones of the lung. In recumbent subjects, the posterior regions are dependent. Lying on the side causes more ventilation to go to whichever lung is lower. This gravity dependence can be exploited to direct ventilation toward healthy lung segments or away from diseased segments by appropriate positioning of the patient.

Time Constants

Compliance and resistance determine local rates of alveolar filling and emptying. This relationship can be shown using the equation of motion where P(t) is set to a constant value, representing a step change (i.e., a sudden change from one level to another, as in from PEEP to inspiratory pressure during pressure control ventilation). For example:

ΔP=PIP−PEEP=V(t)C+RV˙(t)

When this equation is solved for volume as a function of time (using calculus techniques), the result for passive inspiration is:

V(t)=CΔP(1−e−t/RC)

For passive expiration (with any mode of ventilation), it is:

V(t)=CΔP(e−t/RC)

where e is the base of the natural logarithms (approximately 2.72). The product of resistance and compliance (RC in the equation) has units of time (usually seconds) and is called the time constant. It is referred to as a “constant” because for any value of resistance and compliance, the time constant always equals the time necessary for the lungs to fill or empty by 63%. For unit of inspiratory or expiratory time equal to the time constant, lung volume changes by 63%. After two time constants, lung volume has changed 86%; after three time constants, it has changed 95%. This relationship relates to ventilator settings in that for pressure control modes, inspiratory time must be at least three time constants long to deliver 95% of the volume that is possible with the given pressure settings and lung mechanics. For any mode, expiratory time must be set to at least three time constants for the lungs to empty passively to 95% (i.e., 5% of inspired volume still remains) (Figure 10-19).

A lung unit has a long time constant if resistance or compliance is high. Units with long time constants take longer to fill and to empty than units with normal compliance and resistance (see Figure 10-19). Lung units have a short time constant when resistance or compliance is low. Lung units with short time constants fill and empty more rapidly than lung units with normal compliance and resistance.

Time constants affect local distribution of ventilation in the lung. The effects of unequal time constants within the lung are different for volume control (VC) ventilation (with constant inspiratory flow) compared with pressure control (PC) ventilation (with constant inspiratory pressure). For lung units with equal resistance and compliance, both VC and PC result in equal distribution of volume. For lung units with different resistance and compliance but with equal time constants, the distribution of volume depends only on the ratio of resistance or compliance for both modes of ventilation. For lung units with different time constants but equal resistances, VC gives more uniform volumetric expansion and perhaps lower risk of volutrauma than PC. For lung units with different time constants but equal compliances, PC gives more uniform volumetric expansion and possibly lower risk of volutrauma than VC.²²

Frequency Dependence of Compliance

Variations in time constants can affect ventilation throughout the lung. Abnormal ventilation is characteristic of obstruction in the small airways. This type of obstruction occurs in emphysema, asthma, and chronic bronchitis.²³ The time constants of many lung units are increased in obstructive lung disease. These long time constants are mainly caused by increased resistance to flow in the small airways. Loss of normal tissue elastic recoil, such as in emphysema, also contributes to slowed filling and emptying.

At increased breathing rates, units with long time constants fill less and empty more slowly than units with normal compliance and resistance. Increasingly more inspired gas goes to lung units with relatively normal time constants. When more inspired volume goes to a smaller number of lung units, higher transpulmonary pressures must be generated to maintain alveolar ventilation. Compliance of the lung seems to decrease as breathing frequency increases. This phenomenon is called frequency dependence of compliance.⁵ If dynamic compliance decreases as the respiratory rate increases, some lung units must have abnormal time constants. Any stimulus to increase ventilation, such as exercise, may redistribute inspired gas. Mismatching of ventilation and perfusion can result in hypoxemia, severely limiting an individual’s ability to perform daily activities.

Abnormal time constants in lung units and frequency dependence of compliance can have significant effects on patients requiring mechanical ventilation. When ventilation is controlled in terms of volume or flow along with inspiratory-expiratory times, dynamic hyperinflation (air trapping) can result. Lung volume can increase with mechanical ventilation in a manner similar to that occurring during exercise. Increased ventilation (i.e., breathing rates or flows or both) exaggerates the differences between lung units with long or short time constants.

Minute Ventilation

Ventilation is usually expressed in liters per minute of fresh gas entering the lungs. The total volume moving in or out of the lungs per minute is called minute ventilation. Minute ventilation (exhaled) is denoted by , which is calculated as the product of frequency of breathing (f_B) times the expired tidal volume:

V˙E=fB×VT

For a healthy adult breathing 12 breaths/min and having a VT of 500 ml:

V˙E=12×500ml=6000ml/min,or6L/min

Minute ventilation is normally driven by the production of CO₂ and depends on the size of the subject and his or her metabolic rate. values range from 5 to 10 L/min in healthy adult subjects at rest.

Alveolar Ventilation

The efficiency of ventilation depends on the volume of fresh gas reaching the alveoli (V_A). Rearranging the previously presented equation for calculating V_T yields the following:

VA=VT−VD

Alveolar ventilation, , is the product of breathing frequency (f_B) and alveolar volume per breath (V_A):

V˙A=fB×VA

In a healthy adult with a respiratory rate of 12, V_T of 500 ml, and dead space (V_D) of 150 ml, alveolar ventilation is calculated as follows:

V˙A=12×(500ml−150ml)=12×350ml=4200ml/min

Compare this volume with that described for minute ventilation. is always less than because of the effect of dead space.

Dead Space Ventilation

Estimation of wasted ventilation is essential to assess the efficiency of ventilation. Dead space can be subdivided into the following two components: anatomic dead space and alveolar dead space. When these are considered together, they often are referred to as physiologic dead space.

Anatomic Dead Space

The volume of the conducting airways (including the nasopharynx and oropharynx) is called the anatomic dead space, or V_Danat. V_Danat averages approximately 1 ml per pound of ideal body weight (2.2 ml/kg). For a subject who weighs 150 lb (68 kg), V_Danat is approximately 150 ml. V_Danat does not participate in gas exchange because it is rebreathed. During exhalation of a 500-ml tidal breath, the first 150 ml of gas exhaled comes from the V_Danat. The remaining 350 ml is alveolar gas. At the end of exhalation, the airways contain 150 ml of alveolar gas. During the next inhalation, this 150-ml volume is rebreathed. Only approximately 350 ml of fresh gas reaches the alveoli per breath.

The common estimation of V_D based on body weight goes back to a study published in 1955.²⁴ More recent research has shown poor agreement between an individual patient’s measured dead space and dead space estimated by this and other “rule of thumb” equations.²⁵ Dead space to tidal volume ratio (V_D/V_T) can be more accurately estimated for mechanically ventilated adult patients using more data available at the bedside²⁶:

VDVT=0.32+0.0106(PaCO2−PETCO2)+0.003(RR)+0.0015(age)

where PaCO₂ is arterial O₂ tension (mm Hg), P_ETCO₂ is end-tidal CO₂ tension (mm Hg), RR is respiratory rate (breaths/min), and age is in years.

Alveolar Dead Space

In addition to the ventilation wasted on the conducting airways, some alveoli may not participate in gas exchange. These alveoli are ventilated but not perfused with mixed venous blood. Without perfusion, gas exchange cannot occur. Any gas that ventilates unperfused alveoli is also wasted (dead space effect). Some alveoli have ventilation out of proportion to their perfusion (high ratios; see Chapter 11). These alveoli also contribute to the inefficiency of ventilation because ventilation in excess of what is needed to arterialize the blood in an alveolus is wasted.

The volume of gas ventilating unperfused alveoli is called alveolar dead space, or V_Dalv. Significant amounts of V_Dalv are pathologic. V_Dalv is usually related to defects in the pulmonary circulation. A common clinical example of such a defect is a pulmonary embolism. A pulmonary embolus blocks a portion of the pulmonary circulation; this obstructs perfusion to ventilated alveoli, creating alveolar dead space. Alveolar dead space occurs in addition to the anatomic dead space. In a normal upright subject at rest, alveoli at the apices of the lungs have minimal or no perfusion and contribute to the total volume of dead space ventilation.

Physiologic Dead Space

The sum of anatomic and alveolar dead space is called physiologic dead space (V_Dphy):

VDphy=VDanat+VDalv

The total volume of wasted ventilation, or physiologic dead space, equals the sum of the conducting airways and the alveoli that are ventilated but not perfused (Figure 10-20).

Physiologic dead space includes both the normal and the abnormal components of wasted ventilation. V_Dphy is the preferred clinical measure of ventilation efficiency. Measuring V_Dphy more accurately assesses alveolar ventilation:

V˙A=fB×(VT−VDphy)

Or:

V˙A=V˙E−V˙Dphy

Physiologic dead space is measured clinically by using a modified form of the Bohr equation.

Dead Space/Tidal Volume Ratio

In clinical practice, V_Dphy is often expressed as a ratio to V_T. This ratio (V_D/V_T) provides an index of the wasted ventilation (anatomic plus alveolar dead space) per breath. Measurement of the V_D/V_T ratio requires measurement (or estimation) of the arterial CO₂ (PaCO₂) and the mixed expired CO₂ (P_ĒCO₂). PaCO₂ is usually measured by obtaining an arterial blood gas specimen but can be estimated from an end-tidal gas sample (P_ETCO₂). P_ĒCO₂ may be collected in a sampling bag or balloon or estimated by means of capnography (see Chapter 18). The ratio is calculated using a modified form of the Bohr equation, which assumes that there is no CO₂ in inspired gas:

VDVT=(PaCO2−PE¯CO2)PaCO

where PaCO₂ is arterial CO₂ tension and is the average CO₂ tension in exhaled gas.

In a normal adult subject who has a PaCO₂ of 40 mm Hg and an average expired (mixed expired) CO₂ of 28 mm Hg,

VDVT=(40−28)40=0.30

This equation indicates that the normal dead space ratio is about 30%. This equation assumes that all of the CO₂ in expired gas comes from ventilated alveoli. If all lung units contributed CO₂ equally to the expired gas and there was no anatomic dead space, P_ĒCO₂ would equal PaCO₂, and the V_D/V_T ratio would be zero. Because of anatomic and alveolar dead space, the P_ĒCO₂ is always less than PaCO₂. In a healthy adult, physiologic dead space is approximately one-third of the V_T, with a normal range of 0.2 to 0.4. The V_D/V_T ratio normally decreases with exercise. Both V_T and V_D increase with increased ventilation during exertion, but the V_T normally increases to a greater degree; the ratio decreases (in healthy subjects). V_D/V_T increases with diseases that cause significant dead space, such as pulmonary embolism.