During forced vital capacity maneuvers in subjects with expiratory flow limitation, lung volume decreases during expiration both by air flowing out of the lung (i.e., exhaled volume) and by compression of gas within the thorax. As a result, a flow-volume loop generated by using exhaled volume is not representative of the actual flow-volume relationship. We present a novel method to take into account the effects of gas compression on flow and volume in the first second of a forced expiratory maneuver (FEV1). In addition to oral and esophageal pressures, we measured flow and volume simultaneously using a volume-displacement plethysmograph and a pneumotachograph in normal subjects and patients with expiratory flow limitation. Expiratory flow vs. plethysmograph volume signals was used to generate a flow-volume loop. Specialized software was developed to estimate FEV1 corrected for gas compression (NFEV1). We measured reproducibility of NFEV1 in repeated maneuvers within the same session and over a 6-mo interval in patients with chronic obstructive pulmonary disease. Our results demonstrate that NFEV1 significantly correlated with FEV1, peak expiratory flow, lung expiratory resistance, and total lung capacity. During intrasession, maneuvers with the highest and lowest FEV1 showed significant statistical difference in mean FEV1 (P < 0.005), whereas NFEV1 from the same maneuvers were not significantly different from each other (P > 0.05). Furthermore, variability of NFEV1 measurements over 6 mo was <5%. We concluded that our method reliably measures the effect of gas compression on expiratory flow.
- flow-volume loop
- chronic obstructive pulmonary disease
- lung mechanics
maximal expiratory flow and forced expired volume in the first second of a forced maneuver (FEV1) are routinely used for evaluation of respiratory function in clinical practice. A unique relationship exists between maximal expiratory flow, lung volume, and intrathoracic pressure generated during a forced expiratory maneuver. The maximal expiratory flow is determined by the lung elastic recoil pressure, the upstream frictional pressure loss, and the relationship between cross-sectional area and transmural pressure at the choke point (9, 23). However, the lung elastic recoil pressure is the primary determinant of the maximal expiratory flow (21) and is influenced by the absolute lung volume (7). In normal subjects, at lung volumes near total lung capacity (TLC), expiratory flow increases with efforts that are more forceful. However, below 75% of TLC, an increase in effort does not result in an increase in flow, and it may even paradoxically reduce the flow. For lung volumes <75% of TLC, the major determining factor in maximum achievable expiratory flow is the lung volume (21).
In a typical clinical pulmonary function laboratory, flow is measured with a flowmeter at the mouth level, and FEV1 is obtained by integration of this flow. However, when the volume change during a forced maneuver is measured simultaneously by expired volume using a pneumotachograph and by a volume displacement body plethysmograph, the expired volume may differ considerably from the volume change measured with body plethysmograph. This difference in volume change is primarily because of compression of intrathoracic gas. According to the basic laws governing gas flow in a rigid tube, gas with an access to an opening does not undergo dynamic compression unless the particle velocity of gas in the tube exceeds 0.3 of the speed of sound. However, in an airway that is nonrigid, the governing process of flow limitation is the wave speed concept. At high and mid lung volumes, as flow increases to equal that of wave speed, flow limitation will occur regardless of the driving pressure gradient (6, 15, 20). It is at this point that gas compression may occur, even in areas of lung with access to open airways. In contrast, at lower lung volumes, maximal flow is primarily determined by the coupling of viscous pressure losses and airway mechanical properties. Furthermore, dynamic narrowing of airways during a forced expiratory maneuver in subjects with flow limitation can create areas of trapped gas in the lung (20) and result in gas compression.
Figure 1 represents a forced maneuver in a normal subject and a subject with chronic obstructive pulmonary disease (COPD) using simultaneous measurements obtained with pneumotachograph at the mouth level and body plethysmograph. The data in Fig. 1 demonstrate that volume change measured at the mouth level is not a true representation of the absolute change in lung volume. This is consistent with data from other investigators (4, 5). A flow-volume loop of a patient with severe COPD is shown in Fig. 1B. This patient's measured TLC was 10.23 liters, and his predicted TLC was only 6.69 liters. Interestingly, this subject's forced vital capacity (FVC) was only 2.71 liters. At 50% of vital capacity (or 130% of predicted TLC in this subject), expiratory flow was only 0.4 l/s. At this same flow, the volume detected by the body box is only 90% of predicted TLC. The flow-volume loop generated by using plethysmograph shows a flow of 1 l/s higher than mouth flow at 50% of FVC. In contrast, in a subject with normal lung function (Fig. 1A), there is no appreciable difference between exhaled volume and volume measured with the plethysmograph, and any minor gas compression may be of abdominal gas and airway compression. Interestingly, as the subject relaxes at the end of the FVC maneuver and gas compression is reduced, the lung volume (recorded with plethysmography) increases while there is continuing expiratory flow. Therefore, correlating maximal expiratory flow to the volume measured at the mouth level is a distorted portrayal of the relationship between maximal expiratory flow and absolute lung volume in subjects with expiratory flow limitation (10).
The flow-volume relationship is dependent on the magnitude of thoracic gas compression. During a force maneuver, intrathoracic pressure may reach well above 100 cmH2O. This high pressure can cause significant compression of intrathoracic gas in areas of lungs that are upstream of the choke points. For ease of calculation, we assume a barometric pressure of 1,000 cmH2O. Therefore, a 100-cmH2O intrathoracic pressure will compress the alveolar gas by 10%. Assuming that lung volume of a normal adult man at 50% of vital capacity is ∼3 liters, a 0.3-liter reduction in absolute lung volume resulting from gas compression would occur if this volume of gas was trapped. If the slope of the flow-volume loop were about 2 l·s−1·l−1, a difference of ∼0.6 l/s would be generated because of this gas compression. Therefore, given a typical normal flow of 4 l/s, a variability of ∼15% (0.6 l/s) in flow will be seen because of thoracic gas compression. This degree of variability is well above the clinically significant change in volume and flow seen in response to a bronchodilator in subjects with expiratory flow limitation.
Variables affecting gas compression have been studied both in normal subjects and in individuals with expiratory flow limitation. In a study of normal subjects, Jaeger and Otis (12), showed that magnitude of gas compression increased with increasing airway resistance, exhalation flow rate, and lung volume. Large lung volume and expiratory flow limitations are hallmarks of subjects with COPD. Therefore, in subjects with COPD, gas compression may cause significant distortion of the flow-volume relationship and most likely may result in underestimation of the ventilatory capacity (11). In contrast to expired volume, the volume change measured with plethysmograph is an overestimate of expired volume in an FVC maneuver (11).
In this study, we present a novel method, the no-compression method (NCM), to measure the effects of thoracic gas compression on FEV1. This method involves simultaneous measurement of the volume using a volume-displacement body plethysmograph and mouth flow with a pneumotachograph. The NCM allows us to estimate FEV1 corrected for the effect of gas compression. We present a description of the theory underlying the NCM. Furthermore, to examine the effect of gas compression, we tested the NCM in normal, asthmatic, and COPD subjects.
- FEV1 obtained by mouth flow
- FEV1/predicted FEV1
- FEV1 obtained by the NCM
- NFEV1/predicted FEV1
- (NFEV1 − FEV1)/FEV1
- Expiratory lung resistance during quiet breathing
- Esophageal pressure at the peak expiratory flow
- Maximum esophageal pressure during the expiratory maneuver
- Peak expiratory flow
- Transpulmonary pressure
- Difference between the lowest and highest FEV1 values during same session/lowest FEV1 at that session
- Difference between the maneuvers with lowest and highest NFEV1 from same session/lowest NFEV1 at that session
We studied 11 healthy subjects, 10 subjects with asthma, and 65 subjects with moderate to severe COPD resulting from emphysema. The subjects’ anthropometric and lung function data are shown in Table 1. Asthmatic and COPD subjects had diagnoses of illness based on the criteria of the American Thoracic Society (1). All subjects were clinically stable at the time of the study. The asthmatic and COPD subjects had stopped the use of short-acting bronchodilator for ≥8 h and long-acting bronchodilator for >24 h before lung function tests. The protocol for all subjects was approved by the Institutional Review Board, and each subject gave written consent. All COPD subjects had ceased smoking ≥3 mo before the beginning of this evaluation.
Standard spirometric measurements were obtained before the study. Lung mechanics were measured with the subjects seated in an air-conditioned volume-displacement plethysmograph. The frequency response of this plethysmograph is adequate up to 10 Hz, and the box volume measurement is linear up to 25 l/s (19). Volume measurements were obtained by integrating, with respect to time, the pressure difference across a pneumotachograph measured by an MP45 Validyne (Northridge, CA) pressure transducer (±2 cmH2O) located in the wall of the plethysmograph. The plethysmograph flow is calculated by knowing the resistance of the pneumotachograph. The flow signal was then integrated to obtain volume. The characteristics of this type of plethysmograph are described elsewhere (16). Flow at the mouth level was measured by a no. 3 Fleisch pneumotachograph connected to an MP45 Validyne pressure transducer (±2 cmH2O). Esophageal pressure was measured by a 10-cm-long thin latex balloon positioned in the lower one-third of the esophagus at ∼38–45 cm from the nostril. The balloon was connected to a pressure transducer (±352 cmH2O; Statham 131). The balloon was filled with ∼1 ml air. We calculated the PTP as the difference between mouth and esophageal pressure. Placement of the balloon was considered correct if PTP remained constant while subjects made gentle respiratory efforts against a small orifice and oral pressure increased. The flow, PTP, and volume displacement transducers were connected to a Validyne CD19A high-gain carrier demodulator amplifier. The signals were collected digitally with a personal computer (133 MHz Pentium Dell) using a National Instruments (Austin, TX) 12-bit data acquisition board (Lab-PC+; see Ref. 14). The data collection system is described in detail elsewhere (17). Signals were collected at a rate of 100 Hz. During each session, we obtained at least three reproducible forced expiratory maneuvers. For forced maneuvers, each subject was instructed to inspire to TLC and then with maximal possible effort expire to residual volume (RV). Quality control measures as outlined by the American Thoracic Society were used to select appropriate maneuvers (8). Each subject was instructed to expire for at least 6 s. In addition, expiratory and inspiratory lung resistances were measured during quiet breathing using our previous model (18). TLC and RV were measured in all subjects. The Software was written and developed in MATLAB (Natick, MA) to compute the estimated parameters for the described methods (22). We used normative values of Black et al. (2, 3) for lung function parameters.
After the data are collected, computational program in MATLAB environment is used to calculate the NFEV1. The software was developed based on the following assumptions.
First, a plot of the expiratory flow signal vs. the box volume signal in an x-y graph is generated (Fig. 2A). This plot represents the forced expiratory flow-volume loop with units of liters per second on the y-axis and liters on the x-axis. Subsequently, the software inverts the expiratory flow signal between TLC and RV, plots box volume signal (liters) on the x-axis and inverts expiratory flow signal (s/l) on the y-axis, and generates a graph (Fig. 2B). This new graph contains a very steep negative slope at the beginning of the maneuver that reverses to a positive shallow slope during the effort-independent portion (<80% TLC) of the FVC maneuver. Subsequently, the software uses the following algorithm to calculate the computed time (Z). where V̇ is mouth flow, t is the time at which each data point is recorded, n is the index of data arrays that is collected, and Vbox is the volume measured by the body plethysmograph.
This computed time (Z), a time based on volume and flow, is the mouth transit time for increments of box volume. Likewise, mouth transit time is smaller at the start of the forced maneuver (at TLC) than at the end of the maneuver (at RV). As an individual becomes more obstructed near the end of a forced maneuver, the likelihood of gas compression is higher. By summing each computed time point, the software reconstructs a time line that represents volume changes based on expiratory mouth flow and body plethysmograph volume.
After generating the computed time, the software calculates the subject's NFEV1 (Fig. 2C). The backward extrapolation technique is used to determine the start time for the NFEV1 calculation. The software uses the computed time for this determination instead of the standard linear time. The start time is determined by assuming that peak flow had begun since the expiration began at TLC. Volume-time curves are shown in Fig. 2D. These curves are recorded at the mouth and are based on mouth-flow and box-volume measurements, as described. To estimate the magnitude of gas compression, we used the following equation: where DEFV1 is the difference between FEV1 and NFEV1, and NFEV1 and FEV1 are absolute values of forced expired volume (liters) in first second as measured by the NCM and standard method.
Measuring variability and reproducibility of NFEV1.
To compare variability of the NFEV1 with that of the FEV1 within same testing session, we analyzed the baseline FEV1 and NFEV1 data for all subjects with three FVC maneuvers. We calculated the difference between FEV1 of FVC maneuvers with the highest and lowest FEV1 and compared this difference with the difference in NFEV1 for the same maneuvers. FEV1-Dif was defined as mean difference in FEV1 between maneuvers with the highest and the lowest FEV1 divided by the lowest FEV1. NFEV1-Dif is defined as the mean difference of NFEV1 between maneuvers with the highest and lowest FEV1 divided by the lowest NFEV1.
Furthermore, to evaluate reproducibility of NFEV1 measurements over a 6-mo interval, we calculated the NFEV1 for baseline and a 6-mo follow-up in COPD patients randomized to the medical arm of lung volume reduction surgery (LVRS). The Houston Veterans Affairs Medical Center LVRS is a randomized trial with a control arm of medical therapy and an intervention arm of bilateral LVRS in patients with severe emphysema. Individuals enrolled in this study had lung function measurement at baseline and at 6 mo follow-up.
We used STATA version 7 (College Station, TX) for data analysis. Demographic and baseline lung function data were analyzed by descriptive statistics and presented as means ± SD. The relationships between DFEV1 (as an index of magnitude of gas compression) and other parameters of lung function were evaluated by stepwise multiple regression analysis. We used Student's t-test and ANOVA when appropriate to compare FEV1 and NFEV1 values for the same-session comparison and between-session comparisons. A P value <0.05 was considered significant.
Baseline pulmonary function data are shown in Table 1. The mean FEV1 was 45% of predicted values, whereas mean NFEV1 was 59% of predicted values. The DFEV1 was notably larger in COPD and asthma subjects compared with normal subjects. In a multiple linear regression analysis, the DFEV1 correlated significantly with FEV1, Rle, TLC, and PEF (P values <0.0001). Additionally, slow vital capacity was higher than FVC.
Variability of FEV1 and NFEV1 within the same testing session as measured by FEV1-Dif and NFEV1-Dif was 13 + 1.3% and 2.9 + 2.4%, respectively. Student's t-test showed a significant difference between FEV1-Dif and NFEV1-Dif (P < 0.0005). To measure the reproducibility of the NCM between different testing sessions, we analyzed the data from the medical arm of the LVRS in 23 subjects with severe COPD. These data are presented in Table 2. NFEV1 and FEV1 measured at baseline were not significant after 6 mo. Data indicate that NFEV1 and other parameters of lung function are reproducible over different test sessions.
Figure 3 is the graphical presentation of FEV1 and NFEV1 from a normal subject, a subject with asthma, and a subject with COPD. In Fig. 3, there are six graphs each showing a plot of a volume vs. time of an FVC maneuver. Data for a normal subject at baseline and after inhalation of 180 mg albuterol are shown in Fig. 3, A and B, respectively. The plot shows that, in spite of the high intrathoracic pressure, the magnitude of DFEV1 (i.e., magnitude of gas compression) is negligible because of low Rle. This observation is consistent with the known factors affecting gas compression (13). Figure 3, C and D, shows the relationship between volume and time during a forced expiratory maneuver in a subject with asthma. Interestingly, although PEF increased 60%, reduction of Rle to 50% of baseline has reduced the magnitude of gas compression from 8 to only 1%. Furthermore, the intrathoracic pressure did not change appreciably in response to a considerable increase in PEF. This can be explained in part by the reduction of Rle. The data from a patient with COPD before and after bronchodilator inhalation are presented in Fig. 3, E and F. In this subject, although Rle decreased ∼60% from baseline, the magnitude of gas compression only decreased from 80 to 57%.
The effects of gas compression on forced expiratory flow and volume have been studied previously (10, 12, 13). However, these studies did not provide a methodology for the correction of these measurements for gas compression. In the present study, we presented a novel method (NCM) for measuring the effects of gas compression on maximal expiratory flow in normal, asthmatic, and COPD subjects. In subjects with expiratory flow limitation, a forced maneuver results in gas compression and dynamic gas trapping. With this method, we measured the amount of gas compression and dynamically trapped gas during a forced expiratory maneuver. The method provided a measure of FEV1 corrected for the effect of gas compression. The method requires a volume displacement body plethysmograph and specialized software. Our results demonstrated that, in subjects with expiratory flow limitation, the NCM provided accurate information about the relationship between maximal expiratory flow and true lung volume.
Our data and those of others (11–13) show that gas compression can significantly distort the relationship between flow and volume during an FVC maneuver (Fig. 1B). Such distortion is magnified in subjects with expiratory flow limitation like COPD patients.
As shown in Table 2, the NCM provides reproducible measurements. The test variability over time was negligible in patients with moderate to severe COPD. In subjects with expiratory flow limitation, FEV1 is affected by the underlying airway disease and the effect of gas compression. Therefore, variability in efforts used in performing FVC maneuvers at the same testing session or different testing sessions may result in a false increase or decrease in recorded FEV1. This effect has been shown to be the result of gas compression (13). NCM reproducibility in successive measurements over time indicates that NFEV1 is a reliable measurement. NFEV1 is corrected for the effect of gas compression; therefore, compared with FEV1, NFEV1 can generate more reproducible and less variable measurements. This is indicated by comparing the FEV1 and NFEV1 obtained from multiple maneuvers during the same session. Our data showed FEV1-Dif to be significantly larger than NFEV1-Dif (P < 0.0005).
In normal subjects, the lung function data using expired mouth volume were similar to those using NCM. This finding was expected because airflow limitation and high effort are required to cause appreciable gas compression. As is shown in Fig. 1A, the amount of expired air is the same when measured at the mouth and with the plethysmograph. This is typical for healthy young subjects, where the RV is determined by the balance between expiratory muscle effort and the outward recoil of the respiratory system. In addition, because of dynamic compression of airways, air trapping is less likely in normal subjects. In the normal subjects, factors affecting the gas compression, Rle and TLC, were within normal limits (Table 1). Therefore, it is not surprising that NEFV1 and FEV1 were not statistically significantly different in this group.
In asthmatic subjects, NFEV1 was significantly larger than FEV1. In addition, the Rle was higher than in normal subjects. This can explain the larger gas compression in asthmatic subjects (DFEV1 = 24 ± 20%) compared with normal subjects (DFEV1 = 10 ± 16%). Furthermore, the mean PEF in our normal subjects is much higher than PEF in asthmatic and COPD subjects. However, thoracic gas compression estimated with the NCM was negligible. This observation demonstrates the significant effect of Rle on the generation of gas compression and air trapping and the sensitivity of the NCM in estimating the effect of gas compression. The COPD subjects in this study showed the highest Rle and TLC among the study groups. Therefore, not surprisingly, the gas compression was highest in this group. The relationship between Rle and gas compression should be evaluated with care. A normal Rle can be accompanied by large dynamic compression, provided the effort is very large. In addition, some groups of patients can have normal Rle but severe flow limitation.
In summary, we presented a novel method for measuring the effects of gas compression on expiratory flow (NCM). Our data demonstrated that the NCM can be used to reasonably estimate the gas compression during a forced expiratory maneuver. Our data showed that the NCM produced an index of pulmonary function that was reproducible in repeated measure (NFEV1). Furthermore, the NFEV1 is sensitive to factors affecting magnitude of gas compression. However, future work should include the investigation of the sensitivity of this new method to changes in Rle, TLC, and effort in a larger sample of subjects with and without underlying obstructive airway disease.
This study was supported by National Heart, Lung, and Blood Institute Grant HL-072839 (A. M. Boriek) and a Veterans Affairs Medical Research Service Merit Award for lung volume reduction surgery (S. Goodnight-White).
↵† Deceased 13 September 2000.
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- Copyright © 2004 the American Physiological Society