Measurement of Lung Volumes—Spirometry

Spirometry is the process of measuring volumes of air moving in and out of the lungs. A spirometer is the device used to measure these volumes. The spirogram is a graphical volume versus time representation of the lung volumes made by the spirometer. A classic water-seal spirometer is shown in Figure 3-5. Knowledge of the function of the water-seal spirometer helps a

person understand the interrelationships between volumes and capacities. (Modern spirometers are electronic, microprocessor-controlled devices.)

While wearing nose clips to occlude the nostrils, the patient inhales and exhales through a large tube inserted snugly into the mouth (see Figure 3-5). Exhaled air enters the inverted cylinder (bell) suspended by a chain and pulley. This bell fits loosely over a smaller cylinder and floats in a water-filled sleeve. The other end of the chain is attached to a recording pen and counterweight, which exactly balance the weight of the floating bell. The water provides a seal from atmospheric air and allows nearly frictionless movement of the bell. Bell movements cause the pen to move up and down, inscribing reciprocal tracings on recording paper wrapped around a drum rotating at a constant speed, creating a graphic representation of volume (vertical axis) versus time (horizontal axis) (Figure 3-6).

Figure 3-6 reveals that not all of the TLC is accessible to the spirometer. The residual volume (RV) cannot be exhaled, even with the greatest expiratory effort, because the rigid rib cage prevents total lung deflation. RV must be measured indirectly through other techniques (see Chapter 5).

Figure 3-6 illustrates the four lung volumes and four lung capacities. The measurements for each are approximate and vary according to an individual’s height, gender, and age. The descriptions and symbols for the static lung volumes and capacities are shown in Table 3-1. (See also Appendix I.)

The tidal volume (V_T) is the volume of air inhaled or exhaled with each breath. The vital capacity (VC) defines the maximum limits of a single breath—that is, from a maximum-effort inspiration to a maximum-effort expiration. Theoretically, the V_T can increase until it equals the VC, but this never occurs, even with the most strenuous exercise. The functional residual capacity (FRC) is the amount of air in the lungs at the point of ventilatory muscle relaxation, also known as the resting level, or end-tidal exhalation level. Normally, about 40% of the TLC is contained in the lung at this resting FRC level. Expiratory (abdominal) muscle contraction is required to exhale any portion of the FRC, which involves the exhalation of expiratory reserve volume (ERV). The balance point between inward lung recoil and outward chest wall recoil determines the FRC level. Any factor that decreases lung recoil force increases the FRC and decreases the inspiratory capacity (IC). Any factor that increases lung recoil force decreases the FRC and decreases the TLC and VC. (Chapter 5 reviews these interrelationships in more detail.)

During resting V_T breathing, only inspiration requires muscular (diaphragmatic) contraction. The end-tidal level is the point of total relaxation for all ventilatory muscles. Tidal exhalation is passive, a result of passive lung recoil. Tidal exhalation ends when lung and chest wall recoil forces reestablish their equilibrium point.

Hooke’s Law and Elastic Recoil

Over the normal physiological range of tidal breathing, the lung’s expansion in response to increasing pressures conforms to Hooke’s law. Hooke’s law states that an elastic structure changes dimensions in direct proportion to the amount of force applied. Each unit of pressure applied to the lung causes an additional unit of volume to be generated. As the lung volume increases, the recoil force increases. This linear relationship is maintained up to the elastic limit. If the elastic limit is exceeded, lung rupture is imminent.

Hooke’s law oversimplifies the response of the lungs to increasing pressures. Although the P-V curve is relatively linear in the V_T range, a curve that plots the VC against the lung’s distending pressure (P_L) is more S-shaped. The deflation curve follows a different path than the inflation curve, which is explained in the next section.

Static Pressure-Volume Curve

The lung’s P_L (i.e., P_A − P_pl) must be measured to create a static P-V curve for the lung alone (not taking the chest wall into account). The P_pl can be measured with an esophageal balloon, as described previously. P_A is measured at the mouth under static conditions (i.e., when air flow is absent, mouth and alveolar pressures are equal).

An inspiratory P-V curve can be constructed (Figure 3-7) by applying pressure to the trachea, inflating the lung in volume increments in a stepwise fashion from RV to TLC. Each volume increment is held in the lung until all airflow ceases and P_L is recorded. If the lung is deflated in an identical stepwise fashion back to RV, an expiratory P-V curve is constructed. The inflation-deflation curves are traced in a counterclockwise fashion when the lung is inflated by positive pressure applied to the trachea. The curves are nonlinear, partly because the collagen and elastic fibers composing the lung parenchyma are not uniform springs. Elastic fibers are easily stretched, whereas collagen fibers are more resistant to deformation. As the lung volume approaches TLC during inspiration, collagen fibers resist stretching more, causing the slope of the P-V curve to flatten at its upper end, creating an upper inflection point (see Figure 3-7).

Hysteresis and Mechanism of Lung Volume Change

The inflation and deflation limbs of the P-V curve in Figure 3-7 trace different paths, showing that for the same distending pressure, lung volume is greater during deflation than inflation. This phenomenon is called hysteresis, which occurs because the lungs dissipate energy during inflation.⁴ In other words, part of the energy used to inflate the lungs is not recovered during deflation; this means it takes less force to keep the lungs inflated during deflation than it does during inflation. This phenomenon has been experimentally shown: at the same pressure, the lung volume fixed in inflation is lower than the volume fixed in deflation.⁴

The phenomenon of hysteresis is related to the mechanism whereby lung volume changes. The idea that lung inflation and deflation are the result of collective balloon-like changes in 300 million alveolar diameters is overly simplistic. It is now widely understood that the major mechanism of lung expansion and retraction involves the sequential opening and closing (recruitment and derecruitment) of peripheral alveoli.4,5 At the end of a maximally forceful expiration (at RV), numerous alveoli are closed; lung volume increases during the subsequent inspiration mainly because an increasingly greater number of alveoli open up, not because individual alveolar diameters increase. The lungs dissipate energy in the alveolar recruitment process, mostly in overcoming the molecular adhesive forces of surface tension—forces not present during expiration when alveoli are already open.⁴

The lung’s recruitment expansion mechanism has been confirmed by in vivo microscopy of subpleural alveoli in animal models; inflation of the lung from 20% to 80% of TLC increased the number of recruited alveoli but did not increase individual alveolar diameters.⁵ A different investigation yielded similar results; when lung inflation pressure was increased from 10 cm H₂O to 20 cm H₂O and then again from 20 cm H₂O to 30 cm H₂O, alveolar size did not change, but the number of recruited alveoli changed in direct proportion to applied pressure.⁴

The alveolar recruitment-derecruitment mechanism is illustrated in Figure 3-8. At the beginning of inspiration (Figure 3-8, A), numerous alveoli are closed, their walls stuck together. As inspiration continues, these alveoli are forced open and recruited by the increasing P_L (Figure 3-8, B). Recruitment occurs in two phases: in the first phase, from deflation to about 50% of the TLC, lung volume increases secondary to a linear increase in the number of alveoli recruited (i.e., the number recruited is directly proportional to P_L). From 50% to 100% of TLC, the rate of recruitment accelerates, and the number of recruited alveoli increases exponentially. This accelerated recruitment rate is reflected by a steeper slope on the P-V curve, marked by a lower inflection point (LIP) on the inspiratory limb (see Figure 3-7).⁴

It was previously thought that the LIP marked the point at which alveolar recruitment was complete and that if expiratory pressure were not allowed to fall below this point in mechanically ventilated critically ill patients, alveolar closure on expiration would be prevented. It is now known that alveolar recruitment continues beyond the LIP to TLC; the entire P-V curve can be thought of as a recruitment curve.4,6 (See Clinical Focus 3-3 for a discussion of the clinically obtained P-V curve and significance of LIP.)

When the lung deflates, all alveoli are already open; the molecular adhesive forces opposing the recruitment process during inspiration are not present during expiration. During expiration, the lungs contain a greater volume for a given distending force than they contain during inspiration, which explains hysteresis, or the separation of the P-V curve’s inspiratory and expiratory limbs.

Clinical Focus 3-3   What Static and Dynamic Pressure-Volume Curves Tell Us in the Clinical Setting

Static P-V curves (no airflow conditions) are sometimes constructed in mechanically ventilated patients with acute lung injuries with the belief that these curves might be helpful in ventilator management. Beginning at the FRC level, the respiratory therapist progressively inflates the lungs in measured increments, pausing after each inflation to record its corresponding pressure. In healthy lungs, the P-V curve over the V_T range is linear, but in severe lung injury (e.g., acute respiratory distress syndrome [ARDS]), the tidal P-V curve often has a sigmoid shape with a distinct lower inflection point (LIP) and possibly an upper inflection point (UIP), as shown in the figure. Traditionally, the LIP was thought to be the critical pressure at which collapsed alveoli reopened, and the UIP was believed to mark the volume at which alveolar overstretching occurred. The conventional wisdom was to ventilate patients with airway pressures between the two inflection points; the ventilator could be set to maintain positive end-expiratory pressure (PEEP) to keep pressures above the LIP on expiration, and V_T could be limited to keep inspiratory pressure below the UIP. This strategy, sometimes called the “open-lung” approach to mechanical ventilation, was used to protect against ventilator-induced lung injury.

However, pressures measured at the endotracheal tube reflect the summation of many simultaneously occurring events in a highly complex lung composed of 300 million alveoli; alveolar P-V relationships inferred from measurements made at the endotracheal tube introduce uncertainty, limiting the helpfulness of P-V curves in ventilator management. It is now widely accepted that alveolar reopening in acute lung injury continues along the linear middle part of the static P-V curve even above the UIP and that the UIP probably indicates the end of alveolar recruitment, not necessarily the beginning of overdistention.⁶ It is possible that overdistention of some alveoli occurring simultaneously with the recruitment of others might delay the appearance of the UIP, which means overdistention and alveolar injury could occur even at pressures below the UIP.⁶ Regarding the use of the LIP to set PEEP at a level that prevents expiratory alveolar closure, one must keep in mind that PEEP is applied on expiration. The LIP is a feature of the inspiratory P-V curve, more related to alveolar reopening during inspiration than it is to alveolar closure, which occurs during expiration. The LIP appears to be of limited relevance to expiratory alveolar closure and optimal PEEP levels.⁶

What causes the LIP, and what does it mean? The LIP is probably due to the instantaneous reopening of many alveoli with the same threshold opening pressures, as occurs in evenly distributed (homogeneous) lung disease. In unevenly distributed (heterogeneous) lung disease, alveoli with different threshold opening pressures coexist, and the LIP may be absent because alveoli tend to reopen in a more sequential fashion rather than in large numbers all at once.⁶ PEEP tends to be more beneficial in homogeneous lung disease in which there is a widespread presence of nonventilated zones, where large numbers of alveoli can be recruited simultaneously. PEEP is not as effective when lung disease is heterogeneous (characterized by interspersed ventilated and nonventilated zones), where an increase in inspiratory pressure is more likely to overdistend already open alveoli rather than recruit large numbers of alveoli en masse. The P-V curve contour may simply indicate the presence of homogeneous or heterogeneous lung disease and the potential for recruitment, not the optimal level of PEEP needed to prevent alveolar closure.⁶

The V_T P-V curve can be displayed in a breath-to-breath real-time fashion on the monitoring screens of modern mechanical ventilators. However, this dynamically traced P-V curve is affected by airflow resistance and the P-V relationships in the ventilator tubing, humidifier, and endotracheal tube; it cannot provide accurate information about the P-V relationships in the alveoli.²⁸ The ventilator pressure sensor measures pressure at the patient connection part of the ventilator tubing. During inflation, tubing pressure increases rapidly (depending on the ventilator’s flow setting) with little initial volume change as endotracheal tube and airway resistances are overcome. Volume increases more rapidly relative to pressure as elastic resistance is overcome. This transition from frictional to elastic resistance creates an “inflection point” on the lower part of the inspiratory P-V loop, but it does not have the same meaning as the inflection point on the static P-V curve.²⁸

Lung Distensibility: Static Compliance

Compliance is a measure of the lung’s opposition to inflation. The P-V curve (see Figure 3-7) is a compliance curve. Lung compliance (C_L) is defined as the change in lung volume produced by a unit of pressure change and is measured in liters per centimeter of water pressure (L/cm H₂O), shown as follows:

CL=ΔV(liters)ΔP(cmH2O)

Compliance can be conceptualized as distensibility; if the lung is easily distended, it has high compliance. A lung with low compliance is stiff and difficult to inflate. Static C_L can be obtained from the P-V curve in Figure 3-9 by measuring the volume change (vertical axis) produced by a pressure change (horizontal axis). Normal compliance of the lung alone (C_L) is 0.2 L/cm H₂O or 200 mL/cm H₂O.

The opposite of compliance is elastance, defined as the change in pressure required to produce a unit of volume change (cm H₂O/L). Elastance is the mathematical reciprocal of compliance (elastance = 1/compliance). The more elastic the lung, the less its compliance. Elastance can be conceptualized as recoil force; highly elastic lungs are stiff and difficult to inflate and exhibit a high recoil force. By convention, in pulmonary medicine, the elastic characteristics of the lung are quantified in terms of compliance rather than elastance.

Compliance and Lung Volume

The value of C_L depends on the volume at which it is measured. The compliance curve is not linear, as shown in Figure 3-9. Near TLC, lung fibers are stretched and close to their elastic limits; a 500-mL inspiration taken at a point near TLC requires more muscular effort than a 500-mL inspiration beginning at FRC. The slope of the compliance curve is steepest and C_L is greatest at FRC; this is beneficial because an individual normally inhales the V_T from the FRC level where volume changes require little effort. Even during strenuous exercise, people breathe over the lower 70% of the P-V curve, keeping the elastic WOB relatively low. The larger inspired volumes during exercise produce higher lung recoil, which helps provide the force needed to produce high expiratory flow rates during rapid breathing rates.

Compliance (L/cm H₂O) is greater in an adult than an infant because adult lungs are larger and accept more volume for a given pressure. Although the inherent elastic properties of lung tissues may be identical in adult and infant lungs, the infant’s measured C_L is lower. To correct for lung size, compliance should be calculated per unit of lung volume (C_L/vol). This measurement is called specific compliance and is generally referenced to the individual’s FRC (specific compliance = C_L/FRC). Adults and infants have about the same specific C_L.

Figure 3-10 illustrates the effects of different disease types on C_L. Emphysema is characterized by a loss of elastic lung tissue, which means the lungs can be easily distended and have an abnormally low recoil force. Small pressure changes produce large volume changes. Weakened elastic lung recoil changes

the equilibrium point between lung and thoracic recoil forces, which increases the FRC. Pulmonary fibrosis is characterized by high lung recoil forces. The FRC equilibrium point is displaced, this time to a lower point in the TLC than normal.

Nature and Composition of Pulmonary Surfactant

Surfactant molecules in a water medium have extremely weak attractive forces for each other and for the surrounding water molecules. Strong intermolecular attractive forces between water molecules push the neutral surfactant molecules up to the surface, forming a thin molecular film. This surfactant film floats on the surface and shields the water molecules below from contact with air (see Figure 3-11, B). The weak intermolecular attractive forces of the floating surfactant film impart a low surface tension to the solution.

Pulmonary surfactant is a complex substance composed of 90% phospholipid and 10% protein. Dipalmitoyl phosphatidylcholine (DPPC) constitutes about 50% of surfactant’s phospholipid content and is primarily responsible for the surface-tension–lowering properties of surfactant.⁸ DPPC forms a tight monolayer (a single-molecule-thick film) at the air-liquid interface of the alveolus. The more this surface monolayer is compressed (as occurs in smaller alveoli), the more it concentrates the DPPC molecules and lowers the surface tension. Type II cells secrete surfactant phospholipids by discharging lamellar bodies into the liquid film that lines the inner alveolar surface; this process is known as exocytosis. Once secreted, lamellar bodies unravel and transform into a lattice-like structure called tubular myelin (Figure 3-15). Tubular myelin is the immediate precursor of the DPPC surface monolayer; when facilitated by a specialized surfactant protein (described in the following paragraphs), it quickly spreads to form a film one molecule thick on the alveolar air-liquid surface through a process called adsorption.

Normal surfactant contains four specialized proteins that can be divided into two groups: hydrophilic, or water-attracting proteins SP-A and SP-D and hydrophobic, or water-repelling proteins SP-B and SP-C. The two hydrophilic proteins SP-A and SP-D are not as critical for surfactant function as the two extremely hydrophobic surfactant proteins SP-B and SP-C; the presence of SP-B in particular is essential for the formation of tubular myelin and its adsorption on the monolayer film surface.⁸

Adequate functional surfactant must be present in the fetal lung at birth to allow a successful transition to air breathing; otherwise, high surface tension causes widespread alveolar collapse and severe respiratory distress. DPPC is first present in lamellar bodies of the human fetal lung at 24 weeks of gestation.⁸ Avery and Mead⁹ were the first investigators to describe low surfactant levels in premature newborns with respiratory distress syndrome (RDS). It was later confirmed that deficient synthesis and secretion of pulmonary surfactant causes RDS to occur in a premature infant. Specifically, the absence of SP-B is responsible for respiratory distress and gas exchange abnormalities in a premature newborn.⁸

The concentration of DPPC, also known as lecithin, increases during the last 20% of gestation in the full-term fetus. At the same time, the concentration of another phospholipid, sphingomyelin, remains relatively constant.¹⁰ The ratio of lecithin to sphingomyelin (L/S ratio) in amniotic fluid is predictive of fetal lung maturation. An L/S ratio greater than 2:1 indicates fetal lung maturity in nondiabetic pregnancies; the L/S ratio is falsely high in mothers with diabetes.¹⁰ The L/S ratio is generally 1:1 at 31 to 32 weeks of gestation and 2:1 at 35 weeks of gestation.¹⁰ Phosphatidylglycerol (PG) is another phospholipid component of surfactant that is produced quite late in gestation. Its presence in amniotic fluid almost guarantees lung maturity; the risk of RDS is low if PG is present.¹⁰ The L/S ratio and the presence of PG in amniotic fluid continue to be the “gold standard” for determining fetal lung maturity.¹⁰ More recently, the lamellar body count in amniotic fluid has been shown to correlate significantly with the L/S ratio and PG levels. The advantage of the lamellar body count is that it can be performed more quickly and easily and is less costly than phospholipid analysis; in addition, it is less subject to false elevations in diabetes.¹⁰ Maternal administration of steroids may accelerate the production of fetal DPPC.⁸

Physiological Significance of Pulmonary Surfactant

In Figure 3-16, curve C shows that when the alveolar diameter is small in the surfactant-deficient lung, extremely high P_L is required to offset surface-tension forces and prevent alveolar collapse. Therefore, alveoli are quite prone to collapse at low lung volumes, as Laplace’s law predicts. In the presence of pulmonary surfactant, surface tension decreases dramatically as the alveolar volume decreases (curve A); this allows small alveoli to remain inflated at an extremely low P_L.

Pulmonary surfactant is physiologically important because (1) it reduces WOB; (2) it reduces the distending pressure required to keep small alveoli open; and (3) it provides a stabilizing influence, allowing alveoli of different sizes to coexist at the same distending pressure. Pulmonary surfactant also helps reduce the tendency for fluid to leave pulmonary capillaries and enter the interstitial space. In the absence of pulmonary surfactant, high surface-tension–retracting forces create too much negative intrapleural and interstitial pressure, which pulls fluid out of the capillaries. By decreasing alveolar surface tension, pulmonary surfactant helps prevent the development of undue negative interstitial pressure.⁸

Relaxation Pressure-Volume Curves

At FRC, the lungs are stretched outward above their unstressed resting volume, and the thoracic cage is compressed inward, below its unstressed resting volume. The lungs and thorax are coupled by the cohesive seal of the parietal and visceral pleurae. If the pleural seal is broken, exposing the intrapleural space to atmospheric pressure, the lungs recoil inward to their resting minimal air volume, and the thoracic cage recoils outward to its unstressed resting volume (about 75% to 80% of the TLC, or 70% of the VC). The relative strengths of outward thoracic recoil and inward lung recoil determine FRC, which is the resting unstressed volume of the intact lung-thorax system.

Passive relaxation P-V curves can be constructed separately for the lungs and thoracic cage. Figure 3-18, A, shows the relaxation pressures for the lung alone (as if it were outside of the thorax) plotted against TLC. Pressure in the lung is 0 at collapsed resting state (see Figure 3-18, A, 1) and increases to its maximum level as the lung is inflated to TLC, where elastic tissues are stretched maximally (see Figure 3-18, A, 3). At each lung-inflation point in the figure, the airway is sealed, airflow stops, and the positive pressure is the result of passive lung recoil. Figure 3-18, B, shows that if the empty thoracic cage is compressed to the RV level, it generates a strong outward recoil force. This force generates a negative intrathoracic pressure if the thoracic opening is sealed and the compression force is removed, allowing the thorax to recoil freely (see Figure 3-18, B, 1). When the thorax is allowed to recoil outward to the volume it would contain at 70% of the VC, the thorax is at rest, generating neither inward nor outward recoil force (see Figure 3-18, B, 2). Inflation beyond this point expands the thoracic cage above its resting level and generates an inward recoil force. In other words, from 70% of the VC to TLC, the thorax recoils in the same direction as the lungs (see Figure 3-18, B, 3).

Figure 3-19 shows relaxation P-V curves for the lungs and thorax plotted on the same graph along with the coupled lung-thorax system curve. The resting equilibrium point between inward lung recoil force and outward thoracic recoil force defines FRC and normally occurs between 40% and 50% of TLC. This FRC balance point between the two opposing recoil forces is shown in Figure 3-19 as the point where the horizontal distances between 0 pressure and the two recoil curves are identical (distance a is equal to distance b in the figure).

The lung-thorax compliance curve (see solid line in Figure 3-19) is the algebraic sum of the lung (dotted line) and thorax (dashed line) compliance curves. At all points along the lung-thorax curve, its slope is less than the slope of either the lung curve or the thoracic curve, which means that the combined lung-thorax system is less compliant than either of its components alone (i.e., C_LT is less than either C_L or C_T). Put in terms of elastance, combined lung-thorax elastance (the reciprocal of compliance) is greater than either lung or thorax elastance; this helps explain why reciprocals of C_L and C_T must be added to obtain the compliance of the total system. Compliance is a measure of the lung’s distensibility, whereas elastance is a measure of the lung’s resistance to inflation. The thorax and lungs actually represent elastances, not compliances, arranged in series with one another. As with an electrical circuit, total resistance is equal to the sum of the resistances in the series (R_T = R₁ + R₂ + R₃ + …). In this equation, R_T represents total resistance, and R₁ through R₃ represent the individual resistances.

Likewise, lung and thorax elastances (resistance to inflation) can be summed to obtain total elastance (E_LT = E_L + E_T). In this equation, E_LT represents total system elastance, and E_L and E_T represent individual lung and thorax elastances. Because compliance is the reciprocal of elastance, reciprocals of compliance must be added to calculate total compliance. This is shown as follows:

Substituting 1/C for E in equation 1:

Relaxation Pressure-Volume Curves and Disease

A change in lung elasticity alone alters the lung-thorax recoil curve because it changes the equilibrium point between lung and thorax recoil forces. Thus, FRC must change.

If the lung at FRC in Figure 3-19 develops decreased elasticity, the equilibrium between lung recoil and thoracic recoil is disrupted; outward, thoracic recoil predominates, pulling the lungs outward to a larger volume. The thorax stops expanding, and the lung stops enlarging when the stretching lung generates enough recoil force to again balance the oppositely directed thoracic recoil force. This establishes a new equilibrium point at a larger volume than before (i.e., FRC is increased, and V_T is breathed at a higher level in TLC). The lung curve and lung-thorax curve would shift leftward, showing that smaller pressures are required to inflate the lung to a given volume. Pulmonary emphysema, characterized by the destruction of elastic lung tissue, produces the changes just

described. Diseases that increase lung recoil force, such as pulmonary fibrosis, shift lung and lung-thorax curves to the right and reduce FRC because increased lung recoil pulls the thorax down to lower volumes.

Figure 3-20 shows P-V relaxation curves for the lungs and thorax under normal conditions (see Figure 3-20, B) compared with pulmonary emphysema (see Figure 3-20, A) and pulmonary fibrosis (see Figure 3-20, C). The lungs of the patient with emphysema in A have extremely low recoil force; they must be stretched to near TLC before they generate enough recoil to counterbalance the thoracic recoil force and establish FRC. FRC is quite large, mainly because the RV is so large. Even minimal air is increased as a result of the low lung recoil; if these lungs are removed from the chest, their recoil is so weak that they are unable to expel much of their air, and they remain hyperinflated even at rest.

The fibrotic lungs in Figure 3-20, C, recoil quite strongly, even at low lung volumes. Not much inflation is required before their recoil is equal to chest wall recoil. The FRC equilibrium point is established at an abnormally low volume. TLC is reduced because the ventilatory muscles are not strong enough to stretch the strongly recoiling fibrotic lung to greater volumes.

Resistance to Gas Flow

When gas flows through a tube, pressure decreases steadily along the tube from the inlet to the outlet. The pressure drop across the tube (P₁ − P₂) is the driving pressure of the gas flow. A driving pressure is required to overcome flow resistance. Greater frictional resistance causes a greater pressure decrease across the tube. Frictional resistance is present between gas molecules themselves (viscosity) and between gas flow and the tube wall. The greater the cohesive forces between gas molecules, the greater the viscosity of the gas and the greater its opposition to flow. Generally, the smaller the tube radius and the greater the gas viscosity, the greater the resistance. The greater the resistance, the greater the pressure decrease along the tube.

Resistance (R) to gas flow through a tube is expressed as the pressure decrease between the inlet and outlet divided by the flow rate (V˙). This is shown as follows:

R=P1−P2v˙

Airway Resistance

In the lung, R_aw is measured in units of cm H₂O/L/sec. For a given flow rate, a large pressure decrease across the airways indicates the presence of high resistance. Also, the higher the flow rate, the greater the pressure decrease between two points in a given airway. Normal R_aw in adults measured at the FRC level at a constant flow rate of 0.5 L per second is about 0.5 to 1.5 cm H₂O/L/sec; that is, to produce a flow of 1 L per second, a person must generate a pressure difference across the airways (between mouth and alveoli) of 0.5 to 1.5 cm H₂O.

The lung volume markedly affects R_aw. For this reason, R_aw is usually measured at FRC (an easily reproducible volume at which people normally breathe). A special instrument called a body plethysmograph measures P_A, mouth pressure, and flow rate simultaneously to calculate R_aw.

R_aw is greater at low lung volumes than high lung volumes. High volumes near TLC create high P_L gradients and mechanically dilate the airways. Also, high elastic recoil at high lung volumes has a tethering effect on airways, pulling them to larger diameters. Both of these effects are reduced at low lung volumes. Near RV, small, noncartilaginous airways are severely narrowed, greatly increasing R_aw. As RV is approached, some of these small airways collapse because P_pl exceeds intraairway pressure.

Neurogenic factors also affect R_aw at different lung volumes. Stretch receptors in the lung (see Chapter 1) are stimulated at high volumes, causing a reflex decrease in parasympathetic bronchomotor tone, which dilates the bronchi. Conversely, at low volumes, parasympathetic tone is unchecked, and the airways narrow.

Laminar Flow

Flow is laminar when it is organized in discrete streamlines, similar to overlapping cylindrical layers moving at different speeds. The gas flowing in the central layer of the stream moves at the highest velocity, whereas the outside layer contacting the airway wall is stationary; this produces a conical stream front (Figure 3-21) or a parabolic velocity profile. The concentric cylindrical layers of gas slide over one another, creating friction between them as they travel at different velocities. This friction is determined by molecular cohesive forces (unique for each gas type) and is the source of gas viscosity. When flow is laminar, the pressure required to produce a given flow rate through the airways is influenced by gas viscosity but not by the molecular weight or density of the gas.²

Poiseuille’s Law and Airway Resistance

Under laminar flow conditions, the pressure required to produce a given flow rate through a tube is defined by Poiseuille’s law (P=V˙8ln/πr4). In this equation, P represents the pressure gradient across the tube, V˙ represents the gas flow rate, l represents the tube length, n represents gas viscosity, r represents the tube radius, and 8 and π represent constants. This equation states that under laminar flow conditions, the pressure required to drive a given flow through a tube increases if the tube length or gas viscosity increase. More striking is the increased pressure needed to maintain a given flow rate when tube radius decreases. The equation shows that if an airway’s radius decreases to one half of its original size (e.g., by bronchospasm or mucosal edema), 16 times more pressure is required to maintain the original flow through the airway. For practical purposes, the constants in Poiseuille’s law can be lumped together to form one constant (K), simplifying Poiseuille’s law to P=K×V˙/r4. Pressure varies with the fourth power of the tube radius. This means the denominator (r⁴) becomes 16 times smaller if the tube radius is cut in half, which means pressure required to keep flow constant becomes 16 times greater. Of all the factors affecting R_aw, the most profound is the airway radius. In practical terms, a person must exert 16 times more effort (work) to maintain the same flow through airways that constrict to half of their original diameters. Put another way, if muscular weakness or fatigue prevents such increased effort, flow decreases to one sixteenth of its original value, drastically reducing ventilation.

Turbulent Flow

In contrast to laminar flow, turbulent flow is chaotic, with many churning eddy currents (see Figure 3-21). Gas molecules swirl in lateral and forward directions. The entire gas stream moves forward with no velocity difference between its center and outside portions; this lack of difference in flow rates across the tube’s diameter gives turbulent flow a blunt rather than conical velocity profile. Turbulent flow in the lung is much noisier than laminar flow because high-velocity gas molecules strike the airway walls much more frequently. Much energy is required to generate the chaotic, swirling eddy currents of turbulent flow; turbulent flow requires a higher driving pressure than laminar flow to produce a given flow rate. In laminar flow, the pressure gradient is directly proportional to the flow rate; if pressure doubles, the flow rate also doubles. In turbulent flow, the pressure gradient is proportional to the square of the flow rate. Doubling the flow rate requires a fourfold increase in driving pressure. Poiseuille’s law does not apply to turbulent flow.

Turbulent flow is more likely to occur when the flow rate is high and when flow abruptly changes direction, for example, through narrow and sharply branching airways. Reynolds’ number (Re) predicts when the flow regimen will change from laminar to turbulent. Re, a dimensionless number, is derived from values for gas density, viscosity, velocity, and airway diameter. Generally, if Re is greater than 2000, flow is turbulent. In contrast to laminar flow, the pressure required to produce a given flow rate in turbulent flow is influenced by the molecular weight of the gas or its density (greater density requires greater pressure) but not by its viscosity. For this reason, helium, a gas with very low density, is sometimes used to reduce WOB in patients with abnormally turbulent airflow, as occurs in severe airway obstruction (see Clinical Focus 3-8).

Tracheobronchial or Transitional Flow

The type of flow referred to as tracheobronchial or transitional occurs where airways branch and has characteristics of laminar and turbulent flow. One or the other type of flow may predominate, depending on the airway geometry and flow velocity.

Types of Flow in the Lung and Distribution of Airway Resistance

Flow patterns in the lung are highly turbulent in the upper airways, trachea, and major bronchi but never turbulent in smaller bronchioles and distal airspaces. Normal quiet breathing commonly produces instantaneous peak flow rates of 1 L per second, or 60 L per minute, at the mouth. Even during quiet breathing, flow is highly turbulent in the pharynx, larynx, and trachea. As the airways branch, airflow is progressively redirected into more channels until finally, at the terminal bronchiole level, flow is distributed across hundreds of millions of airways; this greatly reduces flow velocity, and flow is laminar. Even during maximal ventilation, flow in the peripheral airways remains laminar. For similar reasons, R_aw is normally much lower in small airways than in large airways (see Chapter 1). The increase in the number of small airways and the increase in their collective cross-sectional area far outweigh the decrease in their individual diameters.

The upper airways, including the larynx, mouth, and nose, normally account for about 50% of total R_aw at FRC; airways less than 3 mm in diameter account for less than 10% of the total R_aw.¹² However, in people who have severe emphysema, chronic bronchitis, or asthma, resistance to flow in the lower airways may exceed upper R_aw

Dynamic Pressure-Volume Curves

Figure 3-22 is a plot of superimposed static and dynamic C_L curves obtained during mechanical positive pressure lung

inflation with the respiratory muscles completely at rest. Lung inflation begins at FRC (point A). The straight, solid line (A-C) is the C_st curve, which is constructed by momentarily interrupting air flow at progressively greater volumes and holding the breath (ventilatory muscles relaxed) until all airflow ceases.

The curved solid line (ABC) plots pressures dynamically in real time as lung volume increases during a single uninterrupted lung inflation, defining the C_dyn curve. The C_dyn curve bows away from the straight C_st line, reflecting the “drag” that frictional airway resistance adds to elastic resistance. If lung inflation is abruptly halted and the breath is held at point B on the dynamic curve, airflow stops, and its frictional resistance ceases to exist; the pressure generated by it disappears. Airway pressure falls to point D on the static pressure curve. Pressure at point D reflects only the elastic recoil forces present at that volume, whereas pressure at point B reflects elastic recoil forces plus flow-resistive forces. Thus, pressure B minus pressure D represents the pressure required to overcome R_aw. The white area in Figure 3-22 reflects static elastic resistance, and the shaded area (inspiratory portion of the loop) represents flow-dependent airway resistance. The area enclosed by the curved dashed line (CEA on the expiratory loop) and the straight static curve (ADC) reflects expiratory airway resistance. Modern positive pressure mechanical lung ventilators can graphically display breath-by-breath P-V loops (compliance curves), allowing respiratory therapists to identify and differentiate elastic and frictional resistances to ventilation.

Peak and Plateau Airway Pressure

The same positive pressure breath illustrated in Figure 3-22 can be graphically represented on a pressure versus time curve (Figure 3-23). As the ventilator inflates the lung, pressure increases to a maximum peak value (see Figure 3-23, point B). This peak pressure (P_peak) corresponds with the pressure represented by point B on the C_dyn curve in Figure 3-22 and reflects the sum of elastic recoil force and frictional airway resistance. If the breath is held momentarily (an inspiratory pause), gas flow stops, and R_aw dissipates, as does the pressure that was required to overcome it; airway pressure falls to a lower level and stays there until the person is allowed to exhale. This plateau pressure (P_plat) (see Figure 3-23, point D) reflects only the elastic recoil force of the lung-thorax system, which corresponds to point D on the static curve in Figure 3-22. The P_peak − P_plat difference reflects only frictional R_aw

P_plat is equal to P_A because it is measured under static conditions after pressures equalize across the airways. P_plat is therefore more relevant than P_peak in assessing the potential for pressure-induced alveolar injury (barotrauma) during positive pressure mechanical ventilation.

Computing Lung-Thorax Compliance and Airway Resistance in Mechanical Ventilation

Respiratory system compliance (lung-thorax compliance [C_LT]) and R_aw can be easily computed on a mechanically ventilated patient. C_LT and R_aw are measures of elastic and flow-resistive opposition to ventilation. By programming the ventilator to create an inspiratory pause, as just described, the respiratory therapist can directly measure P_peak and P_plat (see Figure 3-23). The delivered V_T and the inspiratory flow rate are known values programmed into the ventilator. The C_LT is calculated as follows:

CLT=VT/(Pplat−baselinepressure)

Baseline pressure is the airway pressure at the end of expiration, just before inspiration begins. (In critically ill patients, the ventilator is often programmed to maintain some degree of positive end-expiratory pressure [PEEP] in the airway to keep alveoli above their critical closing pressures.) Changes in the C_LT reflect changes in lung recoil forces.

A mechanically ventilated patient’s R_aw can be computed from the ventilator’s inspiratory flow rate (V˙insp) in L per second, P_peak, and P_plat as follows (assuming a constant flow rate):

Raw=(Ppeak−Pplat)/V˙insp

In this case, the computed R_aw includes the resistances of the ventilator tubing and the endotracheal tube, but changes in measured R_aw are usually a reflection of changes in the patient’s airway diameter.

P_peak during mechanical ventilation reflects both elastic and flow-resistive opposition to lung inflation, whereas P_plat reflects elastic resistance only. P_peak, P_plat, C_LT, and R_aw are clinically useful in distinguishing between factors that increase lung recoil forces and factors that increase flow resistance through the airways.

Expiratory Flow Limitation

Muscular effort determines the peak flow rate that can be generated by a maximal-effort exhalation if one begins to exhale from the TLC level. However, after about 20% of the VC is exhaled (20% of the way from TLC to RV), factors other than effort limit the expiratory flow rate. At some point during exhalation from TLC to RV, the pressure inside the airways falls below the P_pl surrounding the airways; as a result, small airways are compressed and narrowed, limiting the maximal achievable expiratory flow rate.

At the start of a forced exhalation from the TLC level, the thorax contracts, causing the normally subatmospheric P_pl to become positive. As a result, P_A also increases. P_A is the sum of static elastic recoil pressure (P_el) and surrounding P_pl (P_A = P_el + P_pl) as shown in Figure 3-25, A. Frictional resistance to airflow causes the pressure along the airway to decrease steadily from the alveolus to the mouth, where P_ao is 0.

As exhalation continues and approaches RV, elastic lung recoil pressure decreases progressively as the shrinking lung approaches its resting zero recoil level. The progressively lower P_el associated with diminishing lung volumes means the maximum driving pressure that can be generated in the alveolus also diminishes (remember that P_A = P_el + P_pl). As the lungs become smaller, the driving pressure and the maximal achievable expiratory flow rate decrease. This decrease in driving pressure is the main reason why pressures inside the airways eventually drop below the surrounding P_pl as the lung nears RV. The point at which intraairway pressure equals the surrounding P_pl is called the equal pressure point (EPP). Downstream from the EPP, the higher surrounding pleural pressure compresses and

narrows the airways. Normally, this dynamic airway compression occurs only as the lung nears RV, where low elastic recoil force limits the maximum driving pressure that can be generated. In patients with emphysema, airway compression and collapse occur well above the RV level because the elastic recoil force of the lung is abnormally low in the first place.

Effort-Dependent and Effort-Independent Expiratory Flow Rate

If forced exhalation starts from the TLC level, the initial flow rate is effort-dependent—that is, greater efforts produce greater initial flows as shown in Figure 3-25. In Figure 3-25, A, the subject was instructed to inhale to TLC and hold the breath, keeping the glottis open. In Figure 3-25, B, still at TLC, the subject was instructed to exhale with a submaximal effort. In Figure 3-25, C, the subject was instructed to inhale to TLC again and this time to exhale with maximal effort. The high lung volume at the beginning of these exhalations produced a high elastic recoil pressure, which made a high P_A possible. Because P_A is the sum of intrapleural and elastic recoil pressures, the higher the elastic recoil pressure, the greater the difference between P_A and the surrounding P_pl. Therefore, in a forced exhalation beginning from TLC, the P_A pressure head is high enough that pressures down the airway (toward the mouth) never fall below P_pl before gas leaves the thoracic airways. EPP and dynamic airway compression described earlier never occur at high lung volumes near TLC. In Figure 3-25, B and C, greater muscular effort produced greater P_A − P_ao differences and greater flow rates.

In Figure 3-26, the subject was instructed exactly as in Figure 3-25 except that the exhalation maneuvers began at 60% of VC instead of from TLC. At this smaller lung volume, the beginning lung recoil is lower, and the difference between maximum achievable P_A and P_pl is relatively small. During even a submaximal exhalation effort (see Figure 3-26, B), the pressure inside the airway soon falls below P_pl, and EPP is reached well inside the thoracic cavity. A maximal expiratory effort (see Figure 3-26, C) fails to produce a greater flow rate because EPP represents a choke point between the alveolus and the mouth. In this situation, the driving pressure that produces the expiratory flow rate is not P_A − P_ao, as it is in Figure 3-25, B and C, but is instead P_A minus EPP choke point pressure, which is identical for submaximal and maximal efforts in Figure 3-26, B and C. One can think of the airway compression at EPP as a valve that controls the flow rate, similar to the way pinching the middle of a garden hose would control the outflow of water at the hose’s end. Greater muscular effort in Figure 3-26, C, simply produced an increase in P_pl that exactly matched the increase in P_A. Therefore, the flow rate remained constant and was effort-independent.

Isovolume Pressure-Flow Curve

The isovolume pressure-flow curve (Figure 3-27) can be constructed after a person repeatedly inhales as deeply as possible to TLC, each time exhaling the VC with different efforts. Each time the person exhales the VC, greater effort is used, until the last VC is exhaled with maximal effort. The flow rate, volume, and P_pl (using an esophageal balloon) can be measured simultaneously during each VC maneuver. For all VC efforts performed, from weak to maximal efforts, the flow rates and the corresponding P_pl values that occurred exactly at 60% of VC are identified. Each flow rate is plotted against each corresponding P_pl value obtained at 60% of each VC effort (see Figure 3-27). This curve is called an isovolume pressure-flow curve because all flows and pressures are measured at the same volume (60% of VC).

The points on the curve represent the flow rates and pressures produced by the different efforts, all of which are measured precisely at the moment the lung deflates through 60% of the VC. The striking feature of this curve is that flow no longer increases when expiratory efforts cause pleural pressures to increase to greater than 10 cm H₂O. Even maximal effort (60 cm H₂O pleural pressure in Figure 3-27) fails to increase the flow rate. After the flow rate plateau is reached, the flow rate becomes effort-independent. This flow limitation is explained by the EPP phenomenon described in the previous section.

If several other pressure-flow curves are plotted at different lung volumes in exactly the same manner just described, a family of isovolume pressure-flow curves is produced, each at a different lung volume (Figure 3-28, A). At 90% of VC, increasing effort always produces increasing flows (i.e., a flow rate at 90% of the VC is effort-dependent). At lower lung volumes (60%, 30%, and 10% of VC), expiratory flows initially increase with greater efforts, but after a certain point, further effort fails to increase flow. Each lung volume has its own maximum achievable flow rate (V˙max). If each V˙max is plotted against its corresponding lung volume (instead of its corresponding pleural pressure), a flow-volume curve is produced that is identical in shape to the flow-volume curve recorded instantaneously from a single exhaled, forced VC maneuver (see Figure 3-28, B). The FVC flow-volume curve is actually a continuous plot of all the maximum flows achievable at lung volumes ranging from TLC to RV. (The use of flow-volume curves in assessing pulmonary function is discussed in Chapter 5.)

Effect of Compliance

Stiff lungs with low compliance achieve equilibrium with the inflation pressure quickly but at a low volume (Figure 3-29, B). Likewise, stiff lungs empty quickly because of their high elastic recoil. Lungs with low compliance and normal R_aw have short time constants. Four to five time constants are still required for complete inflation or deflation, but the absolute time required is less than for a normally compliant lung.

Effect of Airway Resistance

Figure 3-29, C, illustrates that high frictional resistance leads to long inflation and deflation times. If the same instantaneous inflation pressure is applied to normal lungs and lungs with high R_aw, more time is required for pressures to equalize across the high-resistance airways. Four or five time constants are needed for pressure equalization, but each time constant is relatively long. Given adequate time, the lung with increased R_aw eventually achieves the same volume as the normal lung, if inflation pressures are identical (see Figure 3-29, C). Equally important, lungs with high R_aw require long expiratory (deflation) times. Inadequate expiratory time leads to incomplete lung emptying, air trapping, and hyperinflation of the lungs.

Frequency Dependence of Compliance

Because they account for less than 10% of the total R_aw, small airways (<2 mm in diameter) constitute a “silent zone” in which considerable disease may go undetected by conventional pulmonary function tests. One marker of increased small airway resistance is the frequency dependence of compliance.¹⁵ R_aw is so low in people without disease that C_dyn and C_st are essentially equal at normal breathing rates. Even at 60 breaths per minute (high flow rates), C_dyn and C_st are equal in normal people. Therefore, C_dyn is not normally frequency-dependent, which implies that all ventilating lung units have normal and uniformly distributed resistances and time constants.

When R_aw is not uniformly distributed throughout the lung, different areas fill and empty at different rates. As breathing frequency increases, the difference in filling and emptying rates is exaggerated. As the breathing rate continues to increase, the high-resistance units receive less and less ventilation until eventually most of the V_T enters only normal units. High-resistance units require high driving pressures for a given V_T; measured dynamic compliance decreases. If C_dyn decreases by 25% or more when the breathing rate increases from 20 to 60 breaths per minute, compliance is considered to be frequency-dependent, indicating abnormally high small airway resistance. Frequency dependence of compliance may be present even if conventional R_aw measurements are normal.

Ventilatory Pattern and Work of Breathing

Rapid breathing rates produce high flows, which increase the frictional component of work. Large tidal volumes stretch the lung, increasing its recoil force and the elastic component of

work. In health and disease, the body automatically chooses a breathing rate and V_T that minimize the total WOB.

A patient with chronic obstructive pulmonary disease (COPD) has a high R_aw, especially on expiration, and has weak elastic recoil. Therefore, patients with COPD breathe at low respiratory rates to minimize frictional work and allow more exhalation time. At the same time, these patients can afford to breathe more deeply because of the lung’s weak elastic recoil force. Patients with COPD often must use a considerable amount of muscular energy to exhale through high-resistance airways and to assist the emptying force of weakly recoiling lungs.

A patient with pulmonary fibrosis chooses a rapid, shallow breathing strategy. Elastic recoil is high; therefore, work is minimized by breathing with smaller V_Ts. A normal R_aw allows higher breathing rates. To summarize, high R_aw with normal or increased C_L produces slow, moderately deep breathing, whereas normal R_aw with low compliance produces rapid, shallow breathing.

Muscular Weakness and Fatigue

If the respiratory muscles are already weak, increased WOB may lead to muscle fatigue. Weakness refers to the decreased force-generating capacity of the rested muscle; fatigue refers to the decreased force-generating capacity of actively working muscles, reversible by rest.¹⁸ Respiratory muscle weakness is common in critically ill patients and predisposes them to ventilatory failure. Clinical assessment of respiratory muscle fatigue is important in determining a patient’s need for mechanical ventilation.

Overt diaphragmatic fatigue may result from excessive WOB with or without muscle weakness. Overt fatigue is defined as the inability of the diaphragm to sustain a given force of contraction throughout inspiration.²¹ A measure of the diaphragm’s force of contraction is the transdiaphragmatic pressure (P_di). P_di (pressure across the diaphragm) is measured after the subject swallows a tube with both esophageal (above the diaphragm) and gastric (below the diaphragm) balloons attached. P_di is the difference between esophageal and gastric balloon pressures. Any force that resists lung inflation causes the diaphragm to contract more forcefully, which increases P_di.

A key indicator of diaphragmatic fatigue is the ratio between quiet-breathing P_di and the maximal P_di (P_dimax) a person can generate.²¹ This ratio tells the clinician what percentage of the maximal force-generating capacity the patient uses for each breath. The P_di/P_dimax ratio increases when (1) lung recoil or R_aw increase or (2) muscular weakness (i.e., low P_dimax) is present. The P_di/P_dimax ratio in normal persons breathing quietly is about 0.05. Patients with a P_di/P_dimax ratio of 0.4 or more eventually experience diaphragmatic fatigue.^18,19

The “gold standard” index of diaphragmatic fatigue is the tension-time index of the diaphragm (TT_di), which relates the P_di/P_dimax ratio to the time spent in inspiration (T_i).²⁰ T_i is expressed as a fraction of the total breath cycle time, T_i/T_tot,

or the so-called duty cycle. The product of these two ratios produces the TT_di, shown as follows:

TTdi=Pdi/Pdimax×Ti/Ttot

The greater the diaphragmatic fatigue, the greater the TT_di. TT_di takes into account (1) the percentage of maximum achievable inspiratory force the patient uses for each breath and (2) the length of time the force is applied. Clinical studies reveal that a TT_di of 0.15 is the fatigue threshold or critical TT_di above which healthy subjects cannot persistently maintain their breathing patterns.²⁰ It is widely accepted that TT_di greater than 0.15 is highly predictive of diaphragmatic fatigue.²¹

The disadvantage of TT_di is that it is an invasive procedure involving the swallowing of esophageal and gastric balloons, which makes it impractical for routine clinical assessment. In addition, TT_di focuses on diaphragmatic function alone, and does not take into account the contribution that the rib cage and accessory muscles make to vigorous inspiratory efforts. The diaphragm and rib cage muscles are uncoupled, each able to function independently. In quiet, nonstressful inspiration, only the diaphragm is active, but as inspiratory load increases (e.g., breathing through an endotracheal tube), the rib cage muscles take on an increasingly greater share of the inspiratory work.²¹

An alternative noninvasive estimate of TT_di is the tension-time index of the inspiratory muscles (TT_mus), which reflects the performance of all inspiratory muscles involved in respiratory effort. TT_mus uses easily measured mouth occlusion pressures instead of esophageal and gastric pressures. The noninvasive TT_mus has been validated and shown to correlate highly with the traditional invasive TT_di.21,22 (TT_mus has also been called the pressure-time index [PTI].) TT_mus is calculated as follows:

TTmus=PI/PImax×Ti/Ttot

The negative inspiratory pressures, P_I and P_Imax, are measured at the mouth while the subject breathes through a tube that the clinician can abruptly occlude to make pressure measurements. P_I is the mean occlusion pressure measured during normal breathing, and P_Imax is the maximal occlusion pressure the patient can generate. T_i/T_tot is the fraction of the total respiratory cycle spent in inspiration. P_I is estimated from the negative airway pressure generated during normal breathing 0.1 second after the airway is abruptly occluded (P_0.1). (The greater the value of P_0.1, the greater the TT_mus. The reason P_0.1 increases as muscle strength decreases is that when fatiguing respiratory muscles fail to generate adequate tension, the muscle spindles become excited and stimulate the central nervous system, which increases the nervous drive to breathe.²¹) The equation for estimating P_I (the mean inspiratory occlusion pressure) is as follows:²²

PI=5×P0.1×Ti

This method assumes that P_0.1 is the rate at which pressure increases over the inspiratory time (cm H₂O/sec). Because the pressure generated by the inspiratory muscles does not always increase linearly, this method may overestimate P_I.²¹ However, a given individual’s inspiratory time-pressure waveform is quite stable and repeatable; thus repeated measurements of P_I, P_0.1, and TT_mus for any one individual should be reliable.²³

TT_mus is easily measured at the bedside or in the emergency department with the appropriate technology and correlates extremely well with TT_di in normal subjects, patients with COPD, and patients with restrictive lung diseases.21–23 The fatigue threshold of the respiratory muscles above which patients fail to sustain adequate ventilation lies in a critical TT_mus range of 0.27 to 0.43.²⁴ (A normal TT_mus in healthy control subjects was found in one study to be about 0.05.²¹) Because TT_mus reflects the performance of all inspiratory muscles, it has an advantage over TT_di in assessing patients with high inspiratory loads, relevant especially in patients with COPD and restrictive lung disease. These patients inhale using predominantly rib cage muscles rather than the diaphragm, a pattern exaggerated by high inspiratory workloads.^21,23

An increased use of accessory ventilatory muscles, coupled with rapid, shallow breathing, also signals impending diaphragmatic fatigue.²⁵ Accessory muscle use can be detected by palpating the neck and abdomen for muscle tensing. Fatiguing breathing patterns are characterized by the loss of synchronized outward inspiratory movements of the rib cage and abdomen. Asynchrony occurs when the outward movement of the abdomen lags behind the movement of the rib cage during inspiration. Critically ill patients exhibiting asynchrony have an increased risk of ventilatory failure. A visual sign of overt diaphragmatic fatigue is abdominal paradox, in which the abdomen is sucked inward during inspiration as the rib cage expands outward; this is evidence that the rib cage muscles are predominately being used.²⁵

Because rapid, shallow breathing invariably accompanies respiratory muscle fatigue, the frequency-to-VT (f/V_T) ratio (sometimes called the rapid, shallow breathing index) is widely used in the clinical setting to assess an individual’s ability to sustain adequate spontaneous ventilation. In a landmark study of patients being weaned from mechanical ventilation, patients with f/V_T ratios greater than 100 (V_T is expressed in liters and f in breaths per minute) had a 95% likelihood of developing ventilatory failure, whereas patients with f/V_T ratios less than 100 had an 80% probability of successfully sustaining spontaneous ventilation.²⁶ The f/V_T ratio is an attractive clinical tool for assessing WOB because it is highly predictive of ventilatory failure, is easily measured, and does not require patient effort or cooperation.