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© Springer Nature Singapore Pte Ltd. 2021
J.-X. Zhou et al. (eds.)Respiratory Monitoring in Mechanical

2. Respiratory Mechanics

Jian-Xin Zhou1  , Yan-Lin Yang1, Hong-Liang Li1, Guang-Qiang Chen1, Xuan He1, Xiu-Mei Sun1, Ning Zhu1 and Yu-Mei Wang1

Department of Critical Care Medicine, Beijing Tiantan Hospital, Capital Medical University, Beijing, China


2.1 Pressure and Flow Data Acquisition in Respiratory Mechanics Researches

Accurate measurements of pressure and flow rate are essential in respiratory mechanics researches [1, 2]. Apart from airway pressure displayed on the mechanical ventilator, additional sources of pressure measurements, such as tracheal pressure, gastric pressure, and esophageal pressure, are useful for differentiating the influence of the airway resistance and chest wall elastance on the lung mechanics [37]. Flow measurement is fundamental for lung volume evaluation during mechanical ventilation [1, 2, 8]. In this chapter, frequently used techniques and instruments for pressure and flow data acquisition are introduced.

2.1.1 Pressure Acquisition by Pressure Transducer

A pressure transducer is a device that converts pressure into an analog electrical signal. One of the most commonly used in respiratory mechanics is the strain-gauge based transducer, by which the conversion of pressure into an electrical signal is accomplished by the physical deformation of strain gages incorporated in the pressure transducer [9]. This type of transducer measures the pressure relative to atmospheric pressure. For example, peak airway pressure value of 30 cmH2O at sea-level altitude (atmospheric pressure of 760 cmH2O) indicates the absolute pressure of 793 cmH2O.

The accuracy of pressure measurement by a pressure transducer is dependent on regular calibration. Two parameters have to be considered for calibration, gain and offset [10]. Gain is the slope of calibration curve, while offset indicates the transducer output with zero input. Transducer gain and offset are usually calculated according to transducer specifications. Two kinds of calibration methods can be used for pressure measurements, two-point and one-point calibration. Two-point calibration can be used when transducer specifications are unknown [11]. In case of long-term measurement, such as respiratory mechanics study for several days, transducer gain and offset may drift significantly from their nominal values. In this situation, two-point calibration may be useful to regain transducer specifications. For two-point calibration, a known source of input has to be applied, usually using the water column as a reference and zero (atmospheric pressure) as another source of input.

For a certain pressure transducer, the gain value is usually quite stable with short-term duration, such as several hours, while the offset value is prone to a greater drift. In this situation, one-point calibration with only zeroing the transducer (opening to atmospheric pressure) is employed. One-point calibration is based on the hypothesis that the transducer calibration curve only shifts vertically but the slope remains unchanged. It has been strongly recommended to perform the one-point calibration before starting each epoch of data acquisition.

2.1.2 Flow Acquisition by Fleisch Pneumotachograph

The physical basis of flow measurement is Poiseuille’s law of laminar flow. As gas flow through a pipeline with known resistance, the flow rate can be calculated as:

$$ \mathrm{Ptot}=\mathrm{Paw}+\mathrm{Pmus}={P}_0+\mathrm{Ers}\times V+\mathrm{Rrs}\times \dot{V} $$


where $$ \dot{V} $$ is the gas flow rate, ∆P is the pressure drop along the resistance, r is the radius of the pipeline, n is the gas viscosity, and l is the length of the pipeline.

If the resistance is constant and sufficiently low so as not to impede airflow, and if the gas viscosity is unchanged, the flow rate is directly proportional to the pressure drop along this pipeline [12]. Since Poiseuille’s law is applied under conditions of laminar flow, turbulence may result in the overestimation of flow rate. Thus, the accuracy is improved when the turbulent flow is minimized.

The most frequently used airflow measurement is the Fleisch pneumotachograph, which is the gold standard for flow evaluation in respiratory mechanics research [13]. Fleisch pneumotachograph is composed of a large number of brass capillary tubes in a parallel array. This function provides a small fixed resistance to laminar airflow. Two differential pressure transducers are attached to the pressure collecting ports on both the beginning and ending sides of the tube. As gas passes through the tube, pressure drop develops and is measured.

Because the flow rate varies widely in human and animal spontaneous and mechanically ventilated breathings, a single model would not be suitable to provide optimal measurement in this wide range of flow rates. There are different models of Fleisch pneumotachograph with different inner diameter and length of the capillary tube. The most frequently used models for respiratory mechanics research are type 00 and type 000, with respective inner diameter of 6 and 9 mm and length of 75 mm [13]. The dead space of the pneumotachograph also affects the accuracy of measurement. Studies have shown that in a dead space the instrument of 15 mL is the cutoff value for measurement accuracy. The dead space of type 00 and type 000 Fleisch pneumotachograph is 0.9 and 2.0 mL, respectively.

In order to avoid the condensation of water vapor in the Fleisch pneumotachograph channels, the instrument is incorporated with heating apparatus. The new version of Fleisch pneumotachograph is equipped with on-demanded heating system. Otherwise, Fleisch pneumotachograph may be used with the protection of heated-moisture-exchanger.

According to the equation of Poiseuille’s law, the accuracy of measurement by pneumotachograph is affected by several factors including the gas viscosity, temperature, and humidity. Therefore, the Fleisch pneumotachograph should be calibrated before each epoch of measurements using the same source of gas with the same temperature and humidity [14].

Usually, a large syringe (e.g., 1 L) is used to calibrate pneumotachograph with room air. At this time, the effects of temperature and humidity on the measurement should be considered. Gas in the lungs is at body temperature (37 °C) and fully saturated with water vapor. Measured gas through pneumotachograph is approximately at room temperature and partially saturated with water vapor. Thus, for representing the gas volume in the lungs, gas volume measured under ambient temperature and pressure saturated with water vapor (ATPS) or standard temperature and pressure with dry (ATPD) are needed to convert to the volume under body temperature and pressure saturated with water vapor (BTPS). This conversion is based on Charles’s law, which states that gas volume decreases as the temperature decreases [15, 16]. Table 2.1 shows the conversion factors for volume calculation from ATPS to STPD and BTPS.

Table 2.1

Conversion factors for volume calculation from ATPS to STPD and BTPS

Temperature (°C)

ATPS (mmHg)
































































































































ATPS Ambient temperature and pressure, saturated with water vapor, STPD Standard temperature and pressure, dry, BTPS Body temperature and pressure, saturated with water vapor

2.2 Esophageal/Gastric Pressure Monitoring

Mechanical ventilation is an important support therapy, that has been widely used in the intensive care unit (ICU). Mechanical ventilation can improve oxygenation and maintain ventilation function by replacing or assisting the activities of respiratory muscles. In addition, mechanical ventilation can also win valuable time for patients to treat other diseases. However, with more and more extensive application of mechanical ventilation, it is found that mechanical ventilation itself will cause damage to patients to a certain extent. In addition to varying degrees of influence on the hemodynamics of the patients, mechanical ventilation may also lead to the aggravation of the original lung injury and even the occurrence of new lung injury, named as ventilator-induced lung injury (VILI) [17]. In recent years, with the pathophysiology of VILI being explored, the concept of lung protective ventilation strategy has been proposed [18]. Lung protective ventilation strategies individually are helpful to reduce VILI by monitoring respiratory mechanics (including pressure and flow) at the bedside. The airway pressure (Paw) at the open of ventilator tubing reflects the pressure gradient over the entire respiratory system (lung and chest wall). In some cases, such as obesity, thoracic, or abdominal disease, the Paw cannot reflect actualpressure gradient over the lung because the higher percentage of Paw is used to overcome the elastance of the chest wall. The monitoring of pleural pressure (Ppl) or esophageal pressure (Pes) may help to distinguish the pressure gradients acting on the lung and chest wall.

On the other hand, when the patient has spontaneous breathing, the inspiratory muscles and the ventilator both participate in breathing activity during assisting ventilation mode. And the pressure inflating the lung comes from the pressure generated by the ventilator and the patient’s inspiratory muscles. Nowadays, the role of patients’ breathing efforts during mechanical ventilation is still controversial. Some studies have pointed out that a certain degree of breathing effort can maintain the expansion of the lung tissue in the dependent region and further improve oxygenation and prevent the occurrence of diaphragm disused atrophy [19, 20]. Epidemiological data show that the proportion of clinical application of auxiliary ventilation is gradually increasing, suggesting that doctors pay more and more attention to the role of spontaneous breathing [19]. However, excessive breathing effort also can induce spontaneous lung injury (P-SILI) [21]. And studies found that early using continuous infusion of neuromuscular blockers in severe ARDS patients to remove spontaneous breathing can improve patient survival [22]. Therefore, it is very important to monitor spontaneous breathing efforts and to balance the relationship between mechanical ventilation and spontaneous breathing effort during assisting ventilation. The Monitoring of Pes, intragastric pressure (Pga) and a series of parameters derived from Pes and Pga can be used to quantitatively assess spontaneous breathing efforts. This chapter will briefly introduce Pes and Pga monitoring technology from the aspects of physiological basis, technology points, and potential clinical application.

2.2.1 Physiological Basis

The lungs are located in the thoracic cavity and surrounded by the visceral pleura. The visceral pleura and parietal pleura that is close to the chest wall formed the pleural cavity. The lungs and chest wall are connected through the pleural cavity. As two elastic structures, the lungs and chest wall will deform when pressure is applied and recover the original shape after the pressure is removed. The lung has a tendency to recoil and the chest wall has a reverse trend that prevents the lung from collapsing. When a person is in an upright position during resting, Ppl is a slightly negative value.

Understanding the pressure gradient of different parts of the respiratory system is the basis for learning respiratory mechanics. Figure 2.1 and Table 2.2 show the pressure gradient of different respiratory structures. The pressure gradient over the respiratory system (Prs) is the difference between the alveolar pressure (Palv) and the body surface pressure (Pbs); the pressure gradient over the chest wall (Pcw) is the difference between Ppl and Pbs; the pressure gradient over the lung, named as transpulmonary pressure (PL), is the difference between Palv and Ppl; transdiaphragm Pressure (Pdi) is the difference between abdominal pressure (Pab) and Ppl [23, 24]. When the lungs are at resting condition, the glottis is open and no airway closure exists, Palv is equal to Paw. In the formula, Paw can be instead of Palv to calculate Prs and PL. Learning the pressure gradient between different respiratory structures under physiological conditions is very important for us to understand the changes in respiratory mechanics in the state of disease. Observing the relationship between PL and Pcw can reflect the pattern of movement between the chest wall and lung during breathing. Pdi provides a way to evaluate diaphragm movement and quantify spontaneous breathing efforts.


Fig. 2.1

Schematic drawing of pressures for the respiratory system. Paw Airway pressure, Pbs Body surface pressure, Ppl Pleural pressure, Palv Alveolar pressure, Pab Abdominal pressure

Table 2.2

Pressure gradient of different respiratory structures




Total respiratory system

Trans-Respiratory System Pressure (Prs)

Prs = Palv – Pbs = Paw – Pbs

Chest wall

Trans-Chest Wall Pressure (Pcw)

Pcw = Ppl − Pbs


Transpulmonary Pressure (PL)

PL = Palv – Ppl = Paw − Ppl


Transdiaphragm Pressure (Pdi)

Pdi = Pab − Ppl

Palv Alveolar pressure, Pbs Body surface pressure, Paw Airway pressure, Ppl Pleural pressure, Pab Abdominal pressure

During inspiration, the volume of the lungs and chest wall changes with pressure changing. The driving pressure of the respiratory system is generated by the contraction of respiratory muscles during spontaneous breathing, by the ventilator during controlled ventilation and by both during assisting ventilation. The total pressure (Ptot) of the respiratory system includes the sum of the pressure provided by the ventilator (Paw) and the respiratory muscle pressure (Pmus) of the patient. Ptot is used to overcome the elastance and resistance of the respiratory system. The equation of gas motion (2.1) describes the above relationship:

$$ \mathrm{Ptot}=\mathrm{Paw}={P}_0+\mathrm{Ers}\times V+\mathrm{Rrs}\times \dot{V}={P}_0+\mathrm{Ecw}\times V+{E}_{\mathrm{L}}\times V+\mathrm{Rrs}\times \dot{V} $$


P0 is the Paw at the beginning of inspiration, Ers is the elastance of the respiratory system, Rrs is the airway resistance of the respiratory system, V is the change of volume, and $$ \dot{V} $$is the flow. When ventilation mode is set in volume control ventilation with a constant flow, the V and $$ \dot{V} $$ are set parameters. According to Eq. (2.1), Ers and Rrs can be calculated by the inspiratory and expiratory occlusion [23]. In addition, Eq. (2.1) can be converted into different breathing patterns and conditions.

In a controlled ventilation mode and patients have no breathing effort, Pmus is zero and Ptot is completely provided by the ventilator. The elastance of the respiratory system can be distinguished as the elastance of the lungs and chest wall. Equation (2.1) can be transformed into Eq. (2.2):

$$ P=\raisebox{1ex}{$1$}\!\left/ \!\raisebox{-1ex}{$C$}\right.\times \varDelta V+R\times \upsilon +I\times \sigma $$


In Eq. (2.2), Paw, P0, V, and $$ \dot{V} $$can be continuously monitored by the ventilator. When the inspiratory occlusion is preformed, the flow is 0 and the resistance is also 0. The remaining variables Ecw and EL represent the change of Pcw and PL in unit of volume, respectively. And measuring Ppl or Pes is the best way to distinguish the portion of Paw acting on lung and chest wall.

Spontaneous breathing effort refers to the process of respiratory movement and energy consumption driven by respiratory muscles. In patients with spontaneous breathing efforts, Pmus becomes a very important part of the Ptot that drives breathing movements. Under normal physiological conditions, the inhalation is an active process and the exhalation is a passive process. Diaphragm is the main inspiratory muscle, which is located between the thoracic cavity and abdominal cavity. When diaphragm contracts, the downward movement of the diaphragmatic dome enlarges the chest cavity and expands the lungs and chest wall. The diaphragm relaxes as exhale, and the lungs and chest walls return to their original state. The forces of all respiratory muscles can be reflected by monitoring Ppl or Pes. The Pdi calculated with Pab or Pga can reflect the force of the diaphragm during breathing.

2.2.2 Esophagus Pressure Estimates Pleural Pressure

Ppl is an important quantitative index for describing the respiratory mechanical characteristics, but it can only be obtained by invasive methods which cannot be widely used in clinical. As an alternative method, Pes is rather noninvasive method that can be real-time monitored at the bedside. In 1950s, Buytendijk firstly found that Pes could estimate Ppl [25]. Later, researchers found that although Pes tends to be more positive than Ppl, the change of the Pes and Ppl were similar [26, 27]. The measurement of Pes, instead of Ppl, can improve our understanding of the mechanical properties of the lungs and chest wall. A series of derived parameters also can enhance our further understanding of the pathophysiology of respiratory failure and VILI.

It should be noted that when the person is in the upright position, the lower third of the esophagus is very close to the pleural cavity and the Ppl can be passed to the esophagus accurately. At this position, Pes is close to the adjacent Ppl [28]. However, in the supine position, the weight of the mediastinum and abdominal pressure both increase Pes. Some studies have pointed out that the Pes is over Ppl about 5 cmH2O due to the effects of body position and mediastinum [29]. In addition, the heterogeneity of the gravity gradient in the lungs also influence Pes. Pes may underestimate Ppl in dependent regions and overestimate Ppl in nondependent regions. In addition, factors such as lung and chest wall deformation, pleural effusion, esophageal smooth muscle reactivity, increased abdominal pressure, and characteristics of the esophageal manometers may affect the absolute value of Pes. Although the absolute value of Pes is still controversial, it is generally believed that Pes can represent the average level of Ppl in both animal and human studies and the changes between Pes and Ppl have a good correlation. For making sure Pes to reflect Ppl accurately, the technical aspects of Pes measurement are important. It usually includes the characteristics of the esophageal manometers, placement position, balloon inflating volume, and data interpretation.

2.2.3 Technology

The monitor of Pes mainly depends on placing esophageal manometers. There are mainly three types of esophageal manometer tubes: balloon catheters, fluid-filled catheters, and direct pressure sensors. The most commonly used catheters are balloon catheters. Some esophageal balloon catheters have two balloons that can monitor Pes and Pga at the same time. Some esophageal balloon catheters even have the function of nasal feeding, which can be used as gastric tubes. The Pes can be measured by connecting the end of esophageal balloon catheters to the pressure transductor, such as pressure measuring device, the pressure sensor of the ECG monitor, or the auxiliary pressure port integrated into the ventilator. The accuracy of Pes monitoring can be affected by many factors, including the placement of the balloon catheter, the volume of the balloon, and the data interpretation. Position of Esophageal Balloon Catheter and Balloon Volume

First, insert the esophageal balloon catheter into the stomach through the esophagus by the nose or oral, the balloon is inflated in an approximate air and connected to a pressure transductor. Awake patients can be asked to cooperate with swallowing. Then, judge whether the catheter is in the stomach by Pes waveform. For patients with spontaneous breathing, breathing efforts induce Paw decreasing; meanwhile, the diaphragm contraction induces Pga increasing. For patients without spontaneous breathing, the Pes appearing “jagged” waveform when pressing the abdomen continuously suggested that the balloon is located in the stomach. For catheters with gastric tubes function, the position can be judged by extracting gastric juice. Then, withdraw the catheter into the esophagus slowly. For patients with spontaneous breathing, it can be seen that the Pes waveform changes from a positive waveform to a negative waveform during inspiration. For patients without spontaneous breathing, the Pes waveform is always a positive waveform. But it can be seen that Pes obviously increases when the balloon passes through the cardia of esophagus accompanying cardiac artifacts increasing, then the Pes gradually decreases when withdrawing balloon continually [30]. The depth of the catheter tip from the nose is about 35–45 cm. During the placement of the catheter, it is necessary to observe the phenomenon of cough or sudden rise in Paw, avoiding inserting into the airway by mistake [31].

The material, compliance, and volume of the balloon will affect the accuracy of Pes monitoring. Nowadays, the materials of balloon are mainly polyethylene and rubber, the former is more commonly used. Because balloon has the characteristic of compliance, an over-inflation balloon will overestimate the pressure around the balloon. Conversely, low inflation volumes cannot conduct the surrounding pressure and underestimate the surrounding pressure. Several in vitro experiments have testified proper balloon volume of commercially available catheters, and the results show that the volume-pressure curve of the balloon is “S” shaped. These studies provide data for the use of different types of catheters [32, 33]. In addition, the esophagus wall is also an elastic tissue and it also can affect Pes. The study found that the volume-pressure of balloon in clinical also showed an “S” shape with an intermediate linear section, which represents the proper filling volume of balloon. The slope of this linear section also reflects the elastance of the esophagus wall that can be used to correct Pes [34, 35]. It should be noted that the balloon should be reinflated after each adjustment of the catheter position. Occlusion Test

In order to ensure the accuracy of Pes, it needs to be testified by the occlusion test (Fig. 2.2). For patients with spontaneous breathing, occlusion tests are a very classic way to judge the position of the catheter [36]. Expiratory occlusion induces the breathing effort of patients that causes the Paw and Pes to decrease. If the ratio of the change of Paw and Pes (ΔPES/ΔPaw) is between 0.8 and 1.2, the balloon is considered to be in a suitable position, otherwise the location and volume of balloon should be adjusted. The principle of this method is that breathing effort during expiratory occlusion does not change lung volume and transpulmonary pressure, so the changes of Paw and Pes should be consistent. For patients without spontaneous breathing, a positive pressure occlusion test can be performed by compressing chest wall during expiratory occlusion and compare the positive changes in Pes and Paw (Fig. 2.2). The study has shown that the positive pressure occlusion test has a good consistency with traditional occlusion test [37]. Occlusion test has been validated in different types of patients. The position of the balloon, the volume of the balloon, and the patient’s position and lung volume all can affect occlusion test. Before occlusion test, the balloon of the artificial airway should be routinely checked to prevent the air leakage from affecting the occlusion test. In addition to the occlusion test, some catheters have radiopaque markers, and the correct position of the balloon can be judged by a chest X-ray.


Fig. 2.2

Occlusion test in a spontaneous and a passive patient under mechanical ventilation. Breathing effort during expiratory occlusion does not change lung volume and transpulmonary pressure, so the changes of Paw and Pes should be consistent. During patients with spontaneous breathing (right), the ratio of the change of Paw and Pes (ΔPES/ΔPaw) is between 0.8 and 1.2, the balloon is considered to be in a suitable position and appropriate balloon volume. During patients without spontaneous breathing (left), ΔPES/ΔPaw was obtained by pressing the chest wall. Paw Airway pressure, Pes Esophageal pressure Other Factors

During Pes monitoring, the amplitude of cardiac artifacts also affects Pes results, which even can up to 4–5 cmH2O [38]. The study proposed that the use of cardiac cycles to determine the measurement phase of Pes can improve the accuracy of the occlusion test [39]. Adjusting the position of the catheter can lower interference. If the esophageal contraction occurs (the Pes increases significantly and is not related to breath), the measurement of Pes should be stopped until Pes returns to the baseline level. In addition, pay attention to whether the baseline level of Pes is stable. Pes decreasing gradually as time probably suggests balloon leakage, and the esophageal balloon catheter should be replaced.

2.2.4 Application Transpulmonary Pressure during Passive Ventilation

During passive ventilation, the patients do not exist spontaneous breath and all the pressures that drive the respiratory system come from the ventilator. Paw can reflect the mechanics of the respiratory system, but the pressure acting on the lungs is not identified yet. As Table 2.2 shows, PL is the force that directly expands the lung tissue, which is equal to the difference between Palv and Ppl. Under passive condition, inspiratory and expiratory occlusion (usually 3s) can make the air in airway and alveoli reach balance. At this situation, Paw is considered to be equal to Palv. Ppl can be estimated by Pes. And considering the limitations of Pes monitoring, scholars proposed different mathematical models to improve the accuracy of PL calculation. The three main methods are introduced as follows:

Calculation Method of Transpulmonary Pressure

Direct measurement method: The PL is calculated by the difference between the Paw and the absolute value of Pes during inspiratory and expiratory occlusion, PL = Paw − Pes. Due to the weight of the mediastinal organs and the elastance of the esophagus wall in the supine position, the PL is usually underestimated that may not accurately reflect the conditions of alveoli open and close periodically or the alveoli collapse. Therefore, some scholars estimated the weight of the mediastinum based on the change of Pes in different positions of healthy volunteers [40] and estimated the elastance of esophagus wall based on the change of Pes at different balloon volumes [41]. The study also proposed a simple method to correct Pes by minus 5 cmH2O directly [42]. But this correction method is suitable for different types of patients remains to be further verified.

Release derived method: The studies suggested that although there were differences in the absolute values of Pes and Ppl, the correlation and consistency of the changes of Pes and Ppl with breath had a good agreement [26, 27]. The release derived method is based on the changes of Paw and Pes under the conditions of mechanical ventilation and atmospheric pressure (ATM). First, measure Paw and Pes during end-inspiratory and expiratory occlusion. Then, the ventilator is disconnected at the end-expiratory occlusion and the Paw and Pes are measured after patients expiring the residual gas in the lungs, finally calculate the PL according to the formula: PL = (Paw − Paw at ATM) − (Pes − Pes at ATM) [43]. Regardless of the absolute value of Pes, Ppl at the end-expiration is zero during ATM. The advantage of this method is to avoid the influence of mediastinal weight and esophageal wall elastance. The disadvantage is needing to disconnect ventilator, which is possibly harmful for some critically ill patients.

Elastance derived method: This is another method to calculate PL based on pressure changes, proposed by Gattinoni et al. [44]. Under passive ventilation, the Ers and Ecw can be calculated by Paw and Pes obtained from end-inspiratory and end-expiratory occlusion. The Ers is equal to the sum of the EL and Ecw, Ers = EL + Ecw. Using the proportion of EL in the Ers, the PL at end-inspiration can be calculated by the formula: PL = Paw × EL/Ers. This method assumed that the Ers changes linearly during inspiration. The advantage of this method is also to avoid the effect of mediastinal weight; moreover, it does not need disconnecting the ventilator. The disadvantage is that cannot obtain PL at end-expiration.

Some studies compared the agreement of three methods. The PL at end-inspiration calculated by elastance-derived method and release-derived method had a good agreement [45]. Another study measured different lung regions PL in animal models and corpses [46], it found that PL calculated by direct measurement method better reflected PL in the dependent region, which could be used to guide PEEP setting, and PL calculated by elastance-derived method better reflected PL in the nondependent lung region, which can be used to reduce lung overdistension.

Clinical Application

In controlled ventilation, the purposes of PL are to guide the PEEP setting based on PL at end-expiratory, reduce lung overdistension by limiting PL at end-inspiration and ΔPL and evaluate the patient’s recruitability. The main purpose is to lower the risk of VILI and provide evidence for setting mechanical ventilation parameters and other treatments by the above methods.

There is still much controversy about how to select the “optimal PEEP.” According to pathophysiology, PL at end-expiration reflects the state of lung tissue at end-expiration, the negative value indicates the possibility of alveolar collapse or atelectasis. In the EPVent study, Talmor titrated PEEP to keep PL at end-expiration calculated by direct measurement method at 0–10 cmH2O and limit the PL at end-inspiration to less than 25 cmH2O in ARDS patients. The results showed that the oxygenation and compliance were significantly improved, and the mortality also showed a downward trend, compared to the ARDS network low PEEP table method [44]. However, EPVent-2 study did not yield the same results. This study compared the Pes and ARDS network high PEEP table to guide PEEP setting. The results showed that there was no significant difference in setting PEEP levels between the two groups, meanwhile there were no significant differences in other outcomes [47]. The difference in the responsiveness of PEEP (lung recruitability) may be one of the reasons that resulted in different results of these two studies.

As we all know, recruitment maneuver can open the collapsed alveoli, increase lung volume, and improve oxygenation in patients with ARDS. However, the effect of lung recruitment varies in different patients. That is, not all patients are sensitive to recruitment maneuver, as lung recruitability. Some studies pointed out that patients with lung compliance decreasing are mostly recruitable, while those with chest wall compliance decreasing are mostly unrecruitable [48]. For patients with unrecruitable lung, recruitment maneuver probably induces lung overdistension. Therefore, PL derived from Pes monitoring is essential in assessing the effect and safety of recruitment maneuver.

PEEP can prevent the alveoli opening and closing periodically during breathing and play a role in stabilizing the alveoli. However, PEEP may also cause alveoli overdistension to induce VILI. Compared to airway plateau pressure, PL at end-inspiration is the actual pressure that reflects the state of lung inflation. Grasso believes that setting PEEP based on PL at end-inspiration may be safer and more effective. The study titrated PEEP to make PL at end-inspiratory be close to 25 cmH2O in patients with influenza A (H1N1)-associated ARDS and showed that could improve oxygenation and decrease inappropriate use of ECMO [49]. It is currently believed that limiting PL at end-inspiration to below 20–25 cmH2O can reduce the occurrence of lung overdistension and reduce the risk of VILI [49, 50].

With introducing the concept of driving pressure, it is believed that it may be related to mortality in patients with ARDS. The part of the driving pressure acting on the lung is the transpulmonary driving pressure (ΔPL), that may be the key to affecting patient mortality [51, 52]. Studies suggested that for patients with ARDS, ΔPL less than 10–12 cmH2O may help reduce the risk of VILI [50]. But these conclusions need further research and verification. Quantitative Assessment of Spontaneous Inspiratory Efforts

Under assisting/supporting ventilation, it is important to assess the patient’s spontaneous inspiratory efforts, because the pressure inflating the lung comes from the ventilator and patient’s spontaneous inspiratory efforts. Proper spontaneous inspiratory efforts can improve oxygenation and reduce the occurrence of respiratory muscle atrophy that is helpful to wean as soon as possible. Strongly spontaneous inspiratory efforts may lead to P-SILI. Lacking spontaneous inspiratory effort may cause ventilator-induced diaphragmatic dysfunction (VIDD). Clinicians usually judge the extent of spontaneous inspiratory effort by observing the patient’s breathing pattern and amplitude. However, patients might already have existed severe respiratory muscle fatigue if the above symptoms appear, and the symptoms and signs cannot quantitatively assess spontaneous inspiratory effort. Besides, the parameters displayed by the ventilator also cannot assess the patient’s spontaneous inspiratory effort. A combination of Pes and Pga monitoring can quantitatively assess spontaneous inspiratory efforts, including pressure swings, work of breathing, and pressure time product. These parameters can help to better adjust mechanical ventilation settings and guide weaning, identify patient-ventilator asynchronous and intrinsic PEEP (PEEPi).

Quantitative Signs

Pressure swings. Pressure swings refer to the change of pressure during breathing, which is usually equal to the difference between the pressure at the expiratory flow of zero (Pexp) and the pressure at the inspiratory flow rate of zero (Pinsp). The formula is: ∆P = Pexp − Pinsp. The pressure in the formula can be Pes, Pga, and Pdi. During spontaneous breath, Pes decreased and Pga and Pdi increased. The calculation of the pressure swings is very simple and can be monitored in real-time at the bedside. However, the Pes swings do not take into account the effects of time, the elastance of chest wall, and PEEPi. Therefore, it is difficult to assess the energy expenditure of the respiratory muscles.

Work of Breathing (WOB) and Mechanical Power (MP)

When a force is applied to an object, the object moves in the direction of the force, it is considered as working on the object. Working is always accompanied by a change in energy, the international unit is Joule (J). WOB refers to the change of volume caused by pressure during breathing. The formula is: $$ \mathrm{WOB}=\mathrm{P}\times \mathrm{V}=\underset{0}{\overset{v}{\int }} Pdv $$. WOB is usually expressed as joule per liter (J/L). Normal values in healthy adults at resting are approximately 0.35–0.7 J/L [53]. The Campbell diagram drawn by dynamic Pes and lung volume divided the WOB into three parts: resistance, elastance, and PEEP (Fig. 2.3). The limitations of WOB mainly include two points: first, WOB cannot reflect changes in isovolumic contraction because WOB is calculated based on the change of volume induced by pressure; Second, WOB does not consider the effect of time. If the time is different but the same pressure causes the same volume change, the energy is also different. The main energy of WOB is mechanical energy in mechanical ventilation. Combining WOB with respiratory rate can obtain mechanical power, also known as mechanical power (MP). MP integrates several parameters of pressure, volume, and time, which can well reflect the energy consumption of respiratory muscles. Studies showed that the increase in MP is related to the occurrence of VILI [54, 55].


Fig. 2.3

Schematic diagram of WOB and PTPes. (a) Shows the pressure and velocity waveforms in pressure support ventilation mode. From top to bottom, flow velocity, airway pressure, and esophageal pressure were shown, respectively. The part between the two dotted lines in (a) is the part that overcomes PEEPi. It can be seen that Pes has begun to decline, but the flow velocity has not changed. In the Pes pressure-time curve in (a), the colorful area of PTPes includes three parts: overcoming airway resistance (gray), elastic resistance (yellow), and PEEPi (blue). (b) Shows the pressure–volume curve based on Pes and lung volume. The area of the color area indicates that the WOB also includes three parts: overcoming airway resistance (gray), elastic resistance (yellow), and PEEPi (blue). In addition, the black area at the end of breath in (b) indicates the patient’s active expiratory effort. Paw Airway pressure, Pes Esophageal pressure, PEEPi Intrinsic PEEP, WOB Work of breathing, Ccw Chest wall compliance, FRC Functional residual capacity

Pressure Time Product (PTP)

The pressure time product (PTP) refers to the integral of pressure over time. The formula is$$ Ptp=P\times t=\underset{0}{\overset{t}{\int }} Pdt $$, the unit of PTP is cmH2O × s, which is usually calculated in a minute, expressed as cmH2O × s/min. The normal resting state is 60–150 cmH2O × s/min [56]. Applying different pressures can reflect different respiratory muscle conditions. PTP calculated by Pes (PTPes) is used to assess overall respiratory muscle activity. PTP calculated by Pdi (PTPdi) is used to reflect the diaphragm muscles. PTP can also be divided into three parts: resistance, elastance, and PEEPi (Fig. 2.3). Compared with WOB, PTP integrates time parameters and has a good correlation with energy consumption in a certain range. In addition, PTP also can calculate changes in muscle during isovolumetric contraction which is suitable to evaluate ineffective spontaneous breathing efforts in patients with asynchronous and PEEPi. PTP also can guide weaning and adjust analgesic and sedative drugs.

Clinical Application

Although there is still controversy about whether to maintain spontaneous inspiratory effort in patients with mechanical ventilation, it is clear that excessive or insufficient spontaneous inspiratory effort will lead to serious related complications and increasing the incidence of adverse outcomes. Excessive or prolonged ventilator support may lead to inadequate spontaneous inspiratory effort and inducing VIDD. The occurrence of VIDD is associated with adverse outcomes such as prolonged ICU stay. A certain degree of diaphragm contractility can prevent atrophy [57, 58]. On the other hand, strong spontaneous inspiratory effort induces the increase in PL swings that may increase stress and induce lung overdistension, even induce asynchrony, leading to P-SILI. In addition, strong spontaneous inspiratory efforts also cause Ppl decreasing and gas in the lungs redistributing, that is, air flows from the nondependent region into the dependent area. Decreased Ppl can also increase Ptm of blood vessel, leading to pulmonary edema. When the inspiratory load is excessive, the activity of the expiratory muscles increases significantly. Abdominal muscle contraction can lead to increasing intra-abdominal pressure and Ppl. High Ppl might cause alveoli collapsing and atelectasis during expiration. Pes/Pga monitoring during mechanical ventilation should be strengthened in clinical practice to assess patient’s inspiratory effort.

In recent years, the patient-ventilator asynchrony has caused wide attention. The asynchrony can result in severe hypoxemia in patients, increasing WOB, and even increasing mortality and prolonging ICU stays [59, 60]. From the perspective of pathophysiology, asynchrony may cause an increase in ΔPL and/or an increase in tidal volume during ventilation, which may lead to increasing stress and lung overdistension and further inducing VILI or exacerbating lung injury. The above mechanism may be the root reason for adverse outcomes in patients with asynchrony. Many asynchronies cannot be identified by Paw monitoring alone. Only combining with the Pes can distinguish types of asynchrony. It is important to know the types of asynchrony to guide clinical treatment.

For example, double trigger and reverse trigger are two kinds of asynchrony in clinic. These two types are very similar according to Paw waveform. It is difficult to distinguish them without other equipment. However, the occurrence mechanism and treatment of them are quite different. They can be distinguished when applying Pes waveforms. The double trigger is caused by strong inspiratory effort and the second breath is triggered again after the first breath completion. In a reverse trigger, the first breath is triggered by ventilator and then it triggers diaphragm contracting which generates the second breath. Although the two types of asynchrony can both lead to an increase in ΔPL and tidal volume, the mechanisms are totally different and treatments are also different. The double trigger might require increasing the depth of analgesia and sedation, while the reverse trigger is exactly the opposite. Therefore, no matter from the perspective of the pathophysiological mechanism of the disease, or guiding clinical treatment, it is very important to monitor Pes for monitoring asynchrony.

Patients with mechanical ventilation all need to wean. Studies found that PTPes in patients who failed in weaning was higher than those who succeed in weaning [61]. The increase in PTPes may be related to the increase of load on the respiratory muscles. Monitoring Pes and Pga might be helpful to identify the reason for weaning failure and guide treatment. For example, if an increase in airway resistance leads to PTPes increasing, the clinician needs to give treatments such as suction and bronchodilators; if an increase in elastance resistance leads to PTPes increasing, the clinician needs to consider the possibility of pulmonary edema. In addition, Pes swings can also guide weaning by assessing inspiratory efforts that are superior to the rapid shallow breathing index [62].

For patients with spontaneous breathing, Pes is the most accurate way to measure PEEPi. The main causes of PEEPi include incomplete expiration or restricted airflow. Patient’s inspiratory efforts need to overcome PEEPi before increasing lung volume. PEEPi can lead to asynchrony, fatigue of respiratory muscles, and the increase of WOB, even induce lung overdistension and further affect hemodynamics. Under spontaneous breathing conditions, PEEPi is equal to the change of Pes from the beginning of Pes falling to the beginning of the inspiratory flow. However, it should be noted that when the patient expiratory muscles (abdominal muscles) are active, that is, active expiratory appears, the change of Pes should be subtracted from the change in Pab to reflect the actual PEEPi [63]. Assess Heart–Lung Interactions

Cardiopulmonary interaction refers to the effect on the circulatory system when the pressure of the respiratory system changes. The main pressures affecting cardiopulmonary interactions are Ppl and PL. The effect of increased Ppl on the circulatory system is mainly related to tidal ventilation, which produces a preload effect [64]. The effect of increased PL on the circulatory system is mainly related to lung inflation, which produces an afterload effect. In addition, there are differences in the effects of the Ppl and PL on the circulatory system under different ventilation modes. The Pes monitoring can be used to measure the transmural pressure of the cavity. The transmural pressure of the cavity (Ptm) is equal to the inside pressure minus the surrounding pressure. Ptm represents the pressure of the distending cavity and reflects the state of volume in the cavity.

During mechanical ventilation, especially in high PEEP levels, there might be some limitations to assess right ventricular preload and predict volume responsiveness by right atrial pressure (RAP). Applying high levels of PEEP can induce the increase of Ppl and the increase of RAP after pressure is transmitted to the right atrium. As Ppl increase, the resistance of venous return resistance increases, systemic venous return decreases, and right atrial transmural pressure (RAPtm) decreases. Studies found that pulse variability by Pes correction can better predict volume responsiveness [65]. Conversely, an increase in PL can cause an increase in right ventricular afterload and RAPtm. On the other hand, Ppl can affect left ventricular afterload through left ventricular transmural pressure (LVPtm). Left ventricular afterload is positively correlated with LVPtm and the radius of the left ventricle and negatively correlated with the thickness of the left ventricle. In some cases of heart disease (such as dilated cardiomyopathy), the radius of left ventricular increasing and the thickness of left ventricular decreasing results in a significant increase in left ventricular afterload. During mechanical ventilation, an increase in Pes causes a decrease of LVPtm and left ventricular afterload and an increase of cardiac output. Some studies pointed out that an increase in Ppl can cause a decrease of left ventricular afterload, which can increase left ventricular stroke volume and systolic blood pressure in acute heart failure [66].

The dramatic changes in Ppl during spontaneous breathing effort can be shown as a significant change in pulse amplitude with breathing. During spontaneous breathing efforts, Pes increased significantly from −25 mmHg in the inspiration to 8 mmHg in the expiration that resulted in RAPtm increasing significantly. If a patient’s spontaneous breathing effort is too intense due to some diseases, it may cause acute negative pressure pulmonary edema [67]. Its mechanism is related to changes of PL and Ptm of small interstitial blood vessels, which further lead to pulmonary capillary disease and leakage. In addition, a significant decrease in Ppl also leads to increasing systemic venous return, pulmonary circulation pressure, and left ventricular afterload.

2.3 Diaphragm Electromyography (EMG)

The main task of the respiratory muscles is to maintain ventilation and perform non-ventilatory behaviors. As the dominant respiratory muscles, the force generated by the diaphragm can be assessed with transdiaphragmatic pressure (Pdi), which is usually measured with bilateral phrenic nerve stimulation. In normal conditions, the diaphragm generates forces for ventilatory behaviors comprising only 30% of maximal Pdi, whereas non-ventilatory behaviors (e.g., signs) require forces ~60% of maximal Pdi [68, 69]. Near-maximal forces are only generated during expulsive behaviors (e.g., sneezing, coughing).

On the other hand, neuromotor control of respiratory muscles is reflected in muscle electrical activity, which is recorded with electromyography (EMG). Detected with electrodes, usually, the EMG signals are amplified, filtered, and digitized for data analysis. Robust linear correlation between twitch Pdi and peak diaphragm muscle root mean square (RMS) EMG/suggesting that diaphragm EMG measurements may be used as a surrogate of muscle force generation, permitting assessment of neural respiratory drive continuously in various conditions [70, 71].

In critically ill patients, EMG has proven useful in providing better insight into breathing patterns in the control of breathing. The respiratory impulse originated from the neurons of the brainstem, carried via motor nerves, transmit through neuromuscular junctions, and propagate along the respiratory muscle fiber membranes. Any failure of these sites can result in abnormal EMG signals. Combined with other mechanical function tests or alone, the EMG can access the efficacy of the muscle’s contractile function and sometimes, diagnose neuromuscular pathology.

2.3.1 EMG Equipment

Three pathways are established to detect the electrical signal of the diaphragm: transcutaneous, transesophageal, and intramuscular pathway, by placing the sensors on the skin overlying the area of apposition of the diaphragm to the chest wall, swallowed into the esophagus to measure crural diaphragm EMG, or inserted into the diaphragm directly with needle, wire, or hook electrodes. Of these methods, the electrical activity of the heart during contraction inevitably interferes with the electrical signal of the diaphragm measured with a transcutaneous and transesophageal pathway. Due to the characteristics of invasive and discomfort, the intramuscular pathway is impractical to use in critically ill patients.

2.3.2 Surface Electrodes

Not only the diaphragm, but the electrical activity of the accessory muscles could also be monitored with surface electrodes. The advantages of these methods include noninvasive and simple to use. Depending on the researcher’s knowledge of anatomy, the electrodes are placed over or as close as possible to the target muscles, and no standards are recommended in terms of the electrode design or positioning. Thereafter, the main drawback of the surface electrodes to monitor diaphragm is the unreliability of the acquired signals, contaminated with the signals from irrelevant muscles, and variate with the differences of body composition (e.g., subcutaneous fat) [7274]. It must be pointed out that the surface electrodes measure signals just from the costal diaphragm component.

2.3.3 Esophageal Electrodes

Decades ago, investigators mounted pairs of electrodes outside a catheter and acquire the electrical signals of the crural diaphragm by inserting the catheter into the esophagus [75, 76]. In anatomy, the motor innervation zone (a region with a high density of motor endplates) of the crural diaphragm of adults lies 1–3 cm cephalad to the gastroesophageal junction, and compared with the right, the left side approximately 1 cm cephalad [77]. It is very hard to keep the attachment of the electrodes to the monitored crural diaphragm since the electrode catheter could move up to 8 cm during inspiration [78].

Multipair esophageal electrodes have been developed to optimize the diaphragm-electrode positioning, by mounting a serial of ring electrodes on one catheter, fixed distance with each other. All electrodes signals monitored and paired continuously, and the signals from each paired electrodes caudal and cephalad to the central area of the crural diaphragm will have a similar amplitude and opposite polarity, under the condition of the same distance from the center. The signal-to-noise ratio could be enhanced, and the artifact originating from the diaphragm movement could be reduced by double subtract the signals [79]. Based on the acquired reliable diaphragm EMG signals, in 1999, Sinderby et al. developed a novel ventilatory mode, which provides the proportional pressure support to match the patient’s inspiratory effort, known as the neutrally adjusted ventilatory assist (NAVA) [80].

The limitation of the esophageal electrodes includes mini-invasive, discomfort, and relatively expensive of the commercially available catheter. Abnormal esophageal anatomy makes it difficult to place the catheter and increases the risk of complications. Contrary to the surface electrodes, furthermore, the signals only sample the crural portions of the diaphragm, instead of the representation of the whole diaphragm.

2.3.4 Signal Disturbances

The feasibility and reliability of the acquired signals are important for data analysis and interpretation. A good electrical signal should remain stable in frequency and amplitude; however, usually interfered with the power line frequency and movement. The movement of the electrodes and the change of the pressure on the electrode results in large-amplitude, low-frequency artifacts. Since high-pass filters could filter out most of the motion artifacts, loss of low-frequency power from the EMG signal are inevitable. Various originated noises are usually assumed to have constant power density and are estimated with the signal-to-noise ratio.

2.3.5 Cross-talk Signals

Cross-talk signals refer to the signals originating from muscles other than the muscles being investigated. As mentioned above, the amplitude and frequency parameters of the EMG are strongly affected by the cardiac contraction (10 times the power of the diaphragm EMG with a much lower frequency), and it is hard to separate these two components. There are two techniques developed to reduce the electrocardiography (ECG) contamination: the gating technique, which removes a section (0.4 s) of the EMG signal centered on the QRS complex [81] and the double-subtraction technique, which subtracts an ECG template from the diaphragm EMG signal at each occurrence of the ECG waveform [82].

Apart from ECG, both the esophageal peristalsis and abdominal/intercostal muscle activity contribute to the signal contamination of the diaphragm, all of which should be identified and excluded carefully [83, 84].

2.3.6 Application in the ICU

Continuous monitoring of the electrical activity of the diaphragm is very important for critically ill patients. Diaphragm dysfunction is popular in ICU, attributed to the ventilator over- or under-assist, neuromuscular blocking with drugs, and malnutrition, which in turn contribute to the difficult weaning. The diaphragm EMG provides useful information about phrenic nerve and diaphragm function. What is more important, the signals of the diaphragm electrical activity could be used to trigger the ventilator and adjust the level of ventilatory assistance. This novel ventilator mode, specifically, NAVA, has the great advantage of preserving appropriate inspiratory effort during spontaneous breathing [85], improve the patient-ventilator synchrony [86], and facilitate the liberation from ventilator [87].

It is well known that diaphragm active during inspiration and relaxes during expiration. However, it has been demonstrated that contraction of the diaphragm may persist throughout expiration, suggesting a “tonic” activity of the respiratory muscles, especially in healthy premature and full-term newborns, and patients with decreased end-expiratory lung volume (EELV) [88]. When PEEP is applied, the tonic diaphragm activation decreased and the phasic activation increased. This phenomenon could only be detected by the diaphragm EMG, other than other techniques.

2.3.7 Summary

Similar to the electrocardiogram, the electromyography can be used to assess the level of activation of diaphragm fibers and detect the potential neural and neuromuscular pathology. What is more, the electrical activity of the diaphragm could be used to guide the ventilation proportional to the patient’s effort. However, the complexity of data analysis and interpretation combined with the various artifacts limit the application of this technique, which needs to be dealt with in the future.

2.4 Compliance and Resistance

The impedance of respiratory system can be roughly divided into two categories: elastance and inelastic resistance. The former mainly includes the elastic resistance of the lungs and the elastic resistance of the chest wall, which is 2/3 of the main total resistance during breathing. The inelastic resistance includes frictional resistance and inertance, which accounts for about 1/3 of the total resistance when breathing quietly, of which the frictional resistance of the airway is the main.

In respiratory system, the flow of gas is accomplished by the pressure difference, and the respiratory resistance is overcome by driving pressure response. The motion equation of respiratory system is as follows:

$$ P=\raisebox{1ex}{$1$}\!\left/ \!\raisebox{-1ex}{$C$}\right.\times \varDelta V+R\times \upsilon $$


Among them, elastance is often expressed by the reciprocal of compliance (C) and the displacement distance of the respiratory system is expressed by the change of lung volume (∆V). R represents the airway resistance of the respiratory system, υ represents the gas flow rate, I represents the inertance, and σ represents the acceleration of flow. In general, the inertance of the respiratory system is very small and can be ignored. Formula (2.3) can be simplified to:

$$ E={E}_{\mathrm{lung}}+{E}_{\mathrm{chest}\ \mathrm{wall}} $$


2.4.1 Elastance and compliance

Elasticity is one of the main characteristics of respiratory system. As to each given lung volume, there is a corresponding elastance for volume maintenance. Arithmetically, the elastance equals to the sum of lung elastance and chest wall elastance.

$$ 1/{C}_{\mathrm{rs}}=1/{C}_{\mathrm{lung}}+1/{C}_{\mathrm{chest}\ \mathrm{wall}} $$


Compliance is the reciprocal of elastic resistance and is a commonly used clinical concept. Respiratory system compliance (Crs) is a measure of elasticity or distensibility, which can be classified into two categories, compliance of the lungs and chest wall. We can mutate the equation based on Eq. (2.5). The normal respiratory compliance is in the range of 50–70 mL/cmH2O.

$$ {C}_{\mathrm{dyn}}={V}_{\mathrm{T}}/\left({P}_{\mathrm{end}-\mathrm{insp}}-{P}_{\mathrm{end}-\exp}\right) $$

(2.7) Lung Compliance (Clung)

Clung is defined as the lung volume change per unit transpulmonary pressure (PL) gradient change. The formula is: Clung = ∆V/∆PL (L/cmH2O). Clung was S-shaped on the PV curve (Fig. 2.4). The slope of the middle part is the biggest, that is, the compliance is normal, and the lung has good extensibility. The compliance decreased in the upper part of the S-shaped curve. Collagen fibers play a major role in high volume, which is called the inextensibility of the lung. The incidence of lung injury increased significantly at this stage. In the lower part of the curve, there is a low volume segment with alveolar trapping and a significant decrease in compliance. The elastic resistance of the lung comes from two aspects: the surface tension formed by the interface between the liquid layer and the gas on the alveolar surface, and the elastic retraction force of the elastic fibers of the lung, the former accounts for about 2/3 of the pulmonary elastic resistance, and the latter accounts for about 1/3. On the other hand, the compliance of the lung is variable, and the inspiratory branch and expiratory branch of the PV curve are different, and there is a lag phenomenon (Fig. 2.4). And there is a difference in compliance between quick inhalation and slow inhalation. The variability of Clung may be related to the surface tension of alveoli and the viscosity of lung tissue.


Fig. 2.4

S-shaped PV curve of the lung. The abscissa of the figure is transpulmonary pressure and the ordinate is lung volume. The red dotted line indicates the inhalation process and the yellow dotted line indicates the exhalation process. The two curves do not overlap because of the lag phenomenon. The slope of the curve is Clung

In normal subjects, Clung is referred to be 150 mL/cmH2O [90]. when it comes to pulmonary fibrosis, alveolar edema, or discontinuation of ventilation, the compliance is decreased due to the interference of lung inflation or partial atelectasis. In addition, reduced compliance is also found in patients with increased pulmonary venous pressure. For patients with pulmonary emphysema, compliance increases as elastic tissue alters. Notably, compliance also increases during acute asthma attack due to unclear mechanisms. Several factors have been reported to affect compliance, including lung volume, posture, pulmonary blood volume, age, bronchial smooth muscle tone, and comorbidities. Chest Wall Compliance (Ccw)

Ccw is defined as a change in lung volume per unit change in the pressure gradient between the pleural pressure (Ppl) and body surface pressure. The formula is: Ccw = ∆V/∆Ppl (L/cmH2O). Ccw is usually described as milliliters per centimeters of water (mL/cmH2O) or liter per kilopascal (L/kPa). The reference value of Ccw is approximately 200 mL/cmH2O [91]. When the lung volume is 67% of the total lung capacity, the chest is in its natural position. When the lung volume is less than 67% of the total lung capacity, the elastic retractive force of the chest is outward, which is the power to inhale and the resistance to exhale. On the contrary, when the lung volume is greater than 67% of the total lung capacity, the elastic retractive force of the chest is inward, which becomes the resistance to inhale and the power to exhale. Increased age, obesity, thoracic scars are clear sources for reducing Ccw. Moreover, researches have shown that posture is an affecting factor of Ccw. Compared to the supine position, Ccw is elevated with the seat position and lessened in the prone subjects. Measurement of Respiratory System Compliance in Mechanical Ventilation

According to the measurement procedures, Respiratory system compliance (Crs) is classified as dynamic respiratory system compliance (Cdyn) or static respiratory system compliance (Cst) in mechanical ventilation. Cdyn is measured throughout the regular process of normal rhythmic breathing and is the tidal volume (VT) divided by the end inspiration airway pressure minus end exhalation airway pressure both at no flow point, while Cst is VT divided by the pressure gradient that plateau pressure (Pplat) subduce positive end-expiratory pressure (PEEP) which is measured after the inspiratory and expiratory hold. The formula of Cdyn and Cst is:

$$ {C}_{\mathrm{st}}={V}_{\mathrm{T}}/\left({P}_{\mathrm{plat}}-\mathrm{PEEP}\right), $$


$$ \mathrm{Resistance}=\frac{8\times \mathrm{length}\times \mathrm{viscosity}}{\pi \times {\left(\mathrm{radius}\right)}^4} $$


By using an esophageal balloon catheter, a pressure transducer, and a spirometer, Ppl can be measured. By using Ppl, we can calculate PL which is the pressure acting on the lungs. Lung Compliance (Clung) and Ccw can be obtained as lung volume changes dividing by different pressure gradients, respectively.

2.4.2 Resistance

Inelastic impedance includes frictional resistance and inertance. Frictional resistance is the ratio of pressure difference to velocity, and inertance is the ratio of pressure difference to flow acceleration. The frictional resistance of the respiratory system mainly refers to the friction between the gas molecules and the airway wall when the gas flows through the respiratory tract, also known as airway resistance, which is the most common cause of obstructive pulmonary dysfunction in clinic. The relative displacement of chest and lung tissue will also produce frictional resistance. In addition, the inertance of an object mainly depends on the weight per unit volume (density), the degree of change (displacement), and velocity (acceleration). In general, the inertance of the respiratory system is very small and can be ignored.

2.4.3 Gas Flow and Resistance Laminar Flow

When transferred along a regular, straight, and unbranched tube, laminar flow moves as a series of concentric cylinders, where the peripheral of the cylinder moves relatively slower than the center of the cylinder. As a result, a cone is established in front of the laminar flow. Pressure gradient ∆P and flow rate are calculated as:

$$ \varDelta P=\mathrm{flow}\ \mathrm{rate}\times \mathrm{resistance} $$

$$ \mathrm{Flow}\ \mathrm{rate}=\frac{\varDelta P\times \pi \times {\left(\mathrm{radius}\right)}^4}{8\times \mathrm{length}\times \mathrm{viscosity}}, $$
respectively. Therefore, the resistance is qualified as:

$$ \mathrm{Pressure}\ \mathrm{gradient}=k{\left(\mathrm{flow}\right)}^n $$

(2.10) Turbulent Flow

Turbulent flow occurs in high flow rates, tubular angles, diameter change, or branched tubes, which remains an irregular movement that superposes on the gas movement along the tube. Unlike laminar flow, the front of turbulent flow presents with square rather than cone. As to quantitative calculation, resistance concerning turbulent flow is slightly different from that of laminar flow. That is, the driving pressure of turbulent flow is proportional to the square of the gas flow rate, is proportional to the density of the gas, is theoretically fifth power of the radius, and is independent of the viscosity. When gas flow involves partial or total turbulent flow, the pressure gradient can be calculated as follows, Pressure gradient = k1(flow) + k2(flow)2. The parameter k1 and k2 from the equation contain the components of laminar and turbulent flow, respectively. In healthy human subjects, k1 and k2 are summarized to be 0.24 and 0.03. Additionally, the equation can be simplified as:

$$ \mathrm{Pressure}\ \mathrm{gradient}=k{\left(\mathrm{flow}\right)}^n $$


In this equation, parameter n indicates the nature of flow (1—pure laminar flow, 2—pure turbulent flow, between 1 and 2—complex). The parameters K and n in the equation are summarized to be 0.24 and 1.3 in normal human subjects. Reynolds Number

Reynolds number is applied for indicating the nature of gas flow when the progress of gas movement is through a straight tube. Reynolds number can be quantified as follows: linear gas velocity × tube diameter × gas density/gas viscosity. A Reynolds number greater than 4000 indicates that the gas flow is mainly consisted of turbulent flow, whereas less than 2000 indicates laminar flow. A Reynolds number between 2000 and 4000 indicates complex flow that contains both turbulent and laminar flow. Moreover, the Reynolds number is associated with the Entrance length. Entrance length = 0.03 × tube diameter × reynolds number. The entrance length is defined as the distance for establishing laminar flow. As we can conclude from this equation, gas with a low Reynolds number is less resistant and easier to establish laminar flow.

2.4.4 Characteristics of Airway Resistance

Respiratory resistance is consisting of two components, frictional resistance (including airway resistance and tissue resistance) and inertance. Airway resistance is mainly caused by friction. Given the nature of gas flow is much more complex in pulmonary airway than modulation, it is impractical to simply classify the gas flow into laminar or turbulent. As a result of anatomical characteristics, physically, stable laminar flow cannot be established until the last several levels of airway generation. Consequently, excessive turbulent flow potentially creates more frictional resistance through the transduction of gas flow. Tissue resistance is thought to be originated from the lung and chest wall, especially the chest wall, and referred as the viscous force within the tissues. Tissue resistance is rarely related to end-expiratory pressure or tidal volume. It is reported that the tissue resistance owns a 50% weight of total respiratory resistance in anesthetized healthy subjects. In several clinical circumstances, it is important to contemplate the influence of tissue resistance and differentiate the tissue resistance from the total respiratory resistance. Inertance is another component of respiratory resistance. Inertance is difficult to measure and conventionally believed negligible during normal respiration. However, inertance creates significant impedance during high-frequency ventilation.

2.4.5 Influencing Factors of Airway Resistance

Physical compression is one of the major factors that affect respiratory resistance. Physical compression can be classified into volume-related and flow-related airway collapse. Volume-related airway collapse occurs in patients with reduced lung volume. Assuming other factors stay constant, airway resistance is reverse proportional to the lung volume when lung volume declines. Expiratory airway collapse causes “valve” effect as well as gas trapping, which increases the airway resistance and leads to an increase in residual volume and functional residual capacity [92]. Flow-related airway collapse occurs specifically in airways beyond the 11th generation for the absence of cartilaginous structures during the maximal forced expiration, where the Ppl remains higher than the atmosphere [93]. As gas flow along the airway, there will be an equal-pressure point, where the Ppl is equal to the atmosphere. Hence, small airways affiliated to the equal-pressure point will be unable to overcome the transmural pressure gradient and develop airway collapse.

Muscular contraction affects the airway diameter through neural pathways, humoral regulations, reflection to stimulations, and inflammatory responses [94, 95]. Neural pathways mainly act through the parasympathetic system and noncholinergic parasympathetic nerves. Acetylcholine on M3 muscarinic receptors causes contraction of bronchial smooth muscle, while vasoactive intestinal peptide produces muscular relaxation by promoting the production of nitric oxide. Sympathetic system plays a deficient role in airway muscular contraction based on existing evidence. Nevertheless, β2-adrenergic receptors are highly sensitive to circulating adrenaline. When it comes to a sympathetic stress response, elevated adrenaline concentration makes great contributions to muscular tone. Stimulations also give rise to muscular contraction. Mechanical stimulation, inhalation of water, cold air, chemical aerosol, or medications are underlying causes of bronchoconstriction. Local cellular secretion is another important mechanism of muscular contraction. Pathogens or allergens activate the secretion of cytokines and amplify the inflammatory response. Among patients with hyperresponsive airway, this mechanism effortlessly generates bronchoconstriction or even bronchospasm.

2.5 Auto-PEEP

In 1982 [96], Pepe and Marini first used the term “auto-PEEP,” which was the abbreviation of auto generated positive end-expiratory pressure. Pepe and Marini also described the measurement technique and clinical implications in their study. They found that alveolar pressure could remain positive during the expiratory phase, even the PEEP (positive end-expiratory pressure) was not applied. And they noted that this “auto-PEEP” phenomenon could severely depress cardiac output by increasing intrathoracic pressure. Nowadays, studies in this field have been deeply conducted.

2.5.1 Definition and Terminology

From the Greek word, auto means “self.” During the expiratory phase, the respiratory system cannot fully return to relaxed position before the next inspiration initiates. This results in the pressure gradient, which will drive the end-expiratory flow until the sudden stop of inspiratory forces from patient or ventilator. The total end-expiratory alveolar pressure is the sum of the applied PEEP and auto-PEEP.

Auto-PEEP is also termed as PEEPi (intrinsic PEEP) or occult PEEP, which makes confusion of these terms. Auto-PEEP and intrinsic PEEP can be equal when no PEEP is set [97].

Auto-PEEP can be caused by dynamic hyperinflation (DH). The lung volume should return to the relaxation volume at end-expiration in normal conditions, when the patients have airflow obstruction, the lung volume may exceed predicted FRC (functional residual capacity) [98]. Dynamic hyperinflation and air (or gas) trapping (AP) are not always similar, and short expiratory time can produce dynamic hyperinflation without physically trapping gas. For example, in asthma patients without intubation, inspiratory muscles are still working in the early exhalation, and the glottis braking leads to dynamic hyperinflation. On the other hand, gas trapping can be reversed at modest tidal inspiratory pressures in obesity and ARDS (acute respiratory distress syndrome) patients, which have weak inspiratory efforts [97]. So, auto-PEEP can be differentiated into three types: (1) with DH and AP; (2) with DH and without AP; and (3) without DH.

Flow limitation describes the dynamic condition when the flow is limited and cannot be increased anymore, even by increasing alveolar pressure or decreasing airway-opening pressure, which is always related to small airway collapse.

2.5.2 Causes and Determinants

Auto-PEEP can be generated in several pathophysiological and clinical conditions, such as increased airway resistance (including increased equipment expiratory resistance), short expiratory time, long time constant of the respiratory system, high minute volume, tidal expiratory flow limitation, COPD (chronic pulmonary disease), ARF (acute respiratory failure), and obesity.

In fact, expiratory time is a relative concept, and the proportion of expiratory time in the entire respiratory cycle is the key to lead to auto-PEEP. In other words, the time constant of the respiratory system is the key, rather than the seconds of the expiratory time exactly [99].

Dynamic airway collapse is also considered as the intrinsic factor of auto-PEEP [98]. In ventilated COPD patients, the exacerbation of airflow obstruction is caused by dynamic airway collapse and flow limitation.

2.5.3 Effects and Consequences

The effects on respiratory mechanics, gas exchange, and hemodynamics are similar between the intrinsic PEEP (auto-PEEP) and the extrinsic PEEP.

Auto-PEEP is a common cause of dyspnea and patient-ventilator asynchrony (such as noneffective triggering). Studies also show that auto-PEEP is related to VILI (ventilator-induced lung injury), the increase of WOB (work of breathing), and worsen the efficiency of the respiratory muscles.

As for the hemodynamic consequences, the effects of PEEP have been known for more than 60 years. Auto-PEEP can reduce cardiac output by increasing intrathoracic pressure and it can decrease arterial pressure [96]. The dynamic hyperinflation may also cause bradycardia and vasodilation by autonomic reflexes. Auto-PEEP can be a cause of shock and cardiac arrest. Some studies further indicate that it is a common cause of pulseless electrical activity in patients with positive pressure ventilation. During cardiopulmonary resuscitation, auto-PEEP can prevent the return of spontaneous circulation [100].

2.5.4 Detection and Measurement

The auscultation and clinical symptoms can help to suspect the existence of auto-PEEP, the symptoms including (1) the enlargement of chest circumference; (2) lower-effective ventilation; (3) shock, deterioration of cardiovascular function, increased pulmonary artery wedge pressure that cannot be explained with the circulatory system functions; (4) the abrupt increase of the airway peak pressure during volume-preset ventilation and the abrupt decrease of the tidal volume during pressure-preset ventilation, and (5) the decrease of the plateau pressure that cannot be explained with the decrease of respiratory system compliance.

The auto-PEEP can be detected after a EEO (end-expiratory occlusion), with the abrupt increase of the airway pressure. After several seconds of EEO (at least 2–3 s), the stable pressure can be obtained, which is called the total PEEP. Total PEEP is the sum of external PEEP (preset PEEP) and intrinsic PEEP (auto-PEEP).

The intrathoracic pressure changes can be instead of the measurement of esophageal pressure (Pes), which can help to detect auto-PEEP. When the flow still exists during expiratory, the abrupt decrease in Pes indicates that the patient performs an active inspiratory effort. To trigger the ventilator, the inspiratory effort must overcome and counterbalance the auto-PEEP [101]. So in patients with active inspiratory effort, the measurement of auto-PEEP is shown as Fig. 2.5. Auto-PEEP is measured as the negative deflection of esophageal pressure between the start of inspiratory effort and the start of ventilator insufflation (the point of zero flow).


Fig. 2.5

The measurement of auto-PEEP. In patients with active inspiratory effort, auto-PEEP is measured as the negative deflection of esophageal pressure between the start of inspiratory effort and the start of ventilator insufflation (the point of zero flow)

The flow limitation can be detected by manual compression of the abdomen after EEO [99]. The doctor puts his hand on the abdominal wall of the patient, with the palm on the umbilicus oriented perpendicularly to the axis between the xiphoid process and the pubis. Once the insufflation is finished, the doctor makes firm but gentle compression of the abdomen in an anteroposterior direction throughout the whole expiration. Flow limitation is diagnosed when there is all or part of overlap on the flow-volume loops between the compression and passive expiration. If the compression makes the expiratory flow at any volume be higher than passive conditions, the patient has no flow limitation. This method can increase pleural pressure, so it may be appropriate during the convalescent stage of the disease, considering the safety and validity of the method [102].

2.5.5 Management and Treatment

Auto-PEEP can be eliminated by several methods including (1) changes of ventilator parameters; (2) reduction of the patient’s ventilation requirements; (3) reduction of the expiratory resistance, and (4) use of appropriate external PEEP.

In obstructive patients, the reduction in tidal volume and increase in expiratory time can decrease dynamic hyperinflation. As mentioned before, the proportion of expiratory time in the entire respiratory cycle is the key, so increasing inspiratory flow or reducing breathing frequency could more effectively reduce auto-PEEP.

In 1989, Tobin [103] used the model of the waterfall over a dam to explain the effect of external PEEP (the downstream pressure) on auto-PEEP (the upstream pressure). The downstream pressure is the external PEEP, and the upstream pressure is auto-PEEP. According to this theory, applying external PEEP will unload the burden of the inspiratory muscles and help weaning but it will induce VILI if there is no existence of expiratory flow limitation.

Some studies [99] indicated that flow limitation could be effectively reduced by bronchodilators and the sitting position, and it could be the most effective therapies to reduce auto-PEEP by decreasing airway resistance and time constant of the respiratory system.

2.6 Pressure-Time Curve

2.6.1 Background

Although mechanical ventilation is an important sustaining treatment for patients, especially with acute respiratory disease syndrome, it also can cause ventilator-induced lung injury (VILI) when ventilations result in lung atelectrauma and overdistension. In 2000, the shape of dynamic Pressure–time (Pt) curve analysis during constant flow-volume control ventilation was introduced as a noninvasive and real-time method to identify whether lung exists atelectasis and overdistension during ventilation at the bedside [104].

2.6.2 Physiology

Airway pressure-time (Paw-t curve shows the change of airway pressure (Paw) against time, Paw is expressed in cmH2O and time in second (s) on the ventilator waveform. During volume-controlled ventilation with a constant flow, the volume of gas is increased at a constant rate as time, it can be used as a surrogate for a pressure-volume curve to some extent. When assuming airway resistance is constant with inflation, the slope of Paw-t curve reflects the change of respiratory system elastance (as the reciprocal of compliance respiratory system). A linear Paw-t curve represents that the respiratory system elastance remains unchanged throughout tidal inflation (Fig. 2.6a); A downward concavity Paw-t curve represents that respiratory system elastance decreases throughout tidal inflation and suggests existing tidal recruitment, so the positive end-expiratory pressure (PEEP) need to be increased to open collapse lung for avoiding atelectrauma (Fig. 2.6b); An upward concavity Paw-t curve represents that elastance increases because of appearing overdistention with inflation, suggesting that the PEEP, tidal volume, or both need to be decreased (Fig. 2.6c). To qualitatively evaluate the shape of Paw-t curve, the Paw and time can be fitted by the following equation:


Fig. 2.6

The examples of airway pressure-time (Paw-t) curve. The upper figures show the Paw-t curves and the lower figures show the corresponding flow waveforms. The dotted lines represent the beginning or end of the constant flow. During the constant flow, the Paw-t curve in (a) is linear (stress index = 1.005), the Paw-t curve in (b) is downward concavity (stress index = 0.786), and the Paw-t curve in (c) is upward concavity (stress index = 1.156)

$$ \mathrm{Paw}=a\times {\mathrm{time}}^b+c $$
where a represents the slope of Paw-t curve as time is equal to 1 s; c is the value of Paw as the time of 0 s; and b is a dimensionless parameter reflecting the shape of the Paw-t curve, which is named stress index (SI). If SI value is between 0.9 and 1.1, the Paw-t curve is linear, SI < 0.9 is downward concavity shape, and SI > 1.1 is upward concavity shape [105].

2.6.3 Measurement

The Method of Calculating SI Manually: First, adjust the ventilation mode to volume-controlled ventilation with a constant flow. To avoid the influence of spontaneous inspiratory efforts, it is better to use neuromuscular blockade or increase sedation level to make Ramsay score of 5. Second, collect Paw and flow. It is recommended to record the Paw and flow for at least 5 breaths in a high sampling frequency with no phase lag; Third, identify the constant flow segment. The constant flow segment is usually identified as the fluctuation is within ±3% of the steady value. Meanwhile, to eliminate the influence of on-flow and off-flow transients, the constant segment should be further narrowed for 50 ms; Forth, The Levenberg–Marquardt algorithm can be used to fit mean Paw and time based on the equation and calculate the SI and R2 values [105].

Calculations need to be aborted if any following conditions occur: (1) A constant segment is not found because of noise, artifacts, or air leakage; (2) The length of the identified constant segment is less than one-third of the entire inspiratory phase; and (3) R2 values of the fitting are lower than 0.95 [105].

The Method of Calculating SI by Software: At present, an instrument is mainly used in the field of respiratory monitor, (ICU-Lab KleisTEK, Bari, Italy), can be used to collect Paw and flow in 200 Hz by putting a transducer between the Y-piece of the ventilator circuit and the endotracheal tube, and it also can calculate SI value by automatic procedure (Fig. 2.7). Besides, several types of ventilators can automatically measure SI such as Servo-i or Servo-u (Maquet, Sweden), SV800 (Mindray Co. China), and Luft3 (Leistung, Argentina).


Fig. 2.7

The example of stress index calculated automatically by ICU-Lab. The figure shows the result of automatic calculation of stress index by an instrument that is mainly used in the field of respiratory monitor, ICU-Lab (KleisTEK, Bari, Italy). The red is flow waveform and yellow is pressure waveform in the upper that is selected to calculate stress index. And lower left and lower right are the mean flow-time and pressure-time waveform that are selected to calculate stress index. The line A and F represent the beginning and end of inspiration. The line B and E represent the beginning and end of constant flow. The line C and D represent the true beginning and end of constant segment after the constant flow segment is further narrowed for 50 ms for avoiding the influence of on-flow and off-flow transients. The Paw and time between line C and line D can be fitted by the following equation: Paw = a × timeb + c, and b value can be shown automatically (0.989 in this figure)

The Method of Calculating SI by Inspection: Besides automatic measurements of SI from ventilators, obtaining SI is still inconvenient in clinical practice because it needs specific software to collect data or rather complex calculations. The previous study introduced a simple method by visually inspecting the ventilator screen waveforms to evaluate SI reliably and accurately [106]. First, clinical physicians need to freeze the ventilator waveform when patients are ventilated in volume control ventilation with a constant flow. Second, identify the midpoint of the constant inspiratory flow waveform and then find the corresponding point on the Paw-t waveform. Third, put a ruler on the ventilator screen and make it pass through the point as the reference of the tangent line for Paw-t waveform. Forth, judge the relationship of this ruler and Paw-t waveform by visually inspecting:

(1) If the Paw-t waveform almost coincides with the ruler, the Paw-t curve is linear, indicating a SI value between 0.9 and 1.1; (2) if the two sides of Paw-t waveform are both deviating downward from the ruler, the Paw-t curve is downward concavity, indicating an SI < 0.9; and (3) if the two sides of the Paw-t waveform are both deviating upward from the ruler, the Paw-t curve is upward concavity, indicating an SI > 1.1 (Fig. 2.8).


Fig. 2.8

Schematic of the method for visually inspecting the airway pressure-time (Paw-t) waveform and stress index (SI) classification. First, the midpoint on the constant inspiratory flow (red dot) is identified; second, the corresponding point (red dashed line and red circle) on the Paw-t waveform is confirmed; third, a ruler is put on the ventilator screen to mark the tangent line (red solid line) passing through the middle point (a). The relationship of this tangent line and the Paw-t waveform is visually inspected and classified into three categories. The Paw-t waveform almost coincident with the tangent line is judged as a linear shape (b), indicating an SI value between 0.9 and 1.1. The off-line software measured SI to be 0.98 in this case. When the two sides of the Paw-time waveform both deviate downward from the tangent line, this is categorized as a downward concavity (c), indicating an SI 0.9. The off-line software measured SI to be 0.80 in this case. When the two sides of the Paw-time waveform both deviate upward from the tangent line, this is categorized as an upward concavity (d), indicating an SI 1.1. The off-line software measured SI to be 1.20 in this case

Meanwhile, another study also observed the accuracy of direct visual inspection of SI by ventilators’ screen [107]. They enrolled 30 patients with ARDS and set mechanical ventilation in two different inspiratory flow (40 and 60 L/min). The study found that physicians can distinguish three types of SI correctly. Besides, the lower constant flow (as 40 L/min) could improve the sensitivity of visual inspection.

2.6.4 The Specific Pressure–Time Curves

The Sigmoidal Shape of Paw-t Curves: In some situations, the Paw-t curve shows a sigmoidal shape with an initial downward concavity and a final upward concavity, probably because alveolar is opened at the beginning and alveolar exists overdistension at the end as inflation. Under these circumstances, it would be best to divide into two portions and calculate the SI respectively [105]. It was recommended that adjust PEEP to obtain linear in the first portion and set the tidal volume to obtain linear in the second portion (Fig. 2.9).


Fig. 2.9

The example of the sigmoidal shape of pressure-time curve. The figure shows the pressure-time (Paw-t) and corresponding flow waveform. This Paw-t curve presents a sigmoidal shape with an initial downward concavity (a) and upward concavity (b)

The Transpulmonary Pressure-Time Curves: When the esophageal pressure (Pes) was measured by an esophageal catheter, the transpulmonary pressure (PL) can be calculated by the equation as PL = Paw − Pes. The PL and time are also described as the following equation:

$$ \mathrm{PL}=a\times {\mathrm{time}}^b+c. $$

In fact, the previous studies showed the SI obtained by Paw-t curves could reflect SI of obtained PL-t curve well.

2.6.5 Clinical Application of SI

The measurement of SI monitoring provides a noninvasive, repeatable, and real-time method to monitor respiratory mechanics at the bedside.

In 2000, Ranieri et al. first tested PL-t curves profile in an isolated rat lung injury model [104]. They adjusted tidal volume and PEEP to obtain three types of SI. The study found that set a tidal volume as 6–8 mL/kg and adjust PEEP to keep SI within 0.9–1.1 contributed to reducing VILI and had the lowest histological injury score (Il-6 and MIP-2 levels). And further animal studies also found that the SI could identify tidal recruitment and overdistension accurately compared to CT scan [109112].

The SI also has been validated in clinical settings [113117]. Grasso et al. performed a cohort study in 15 ARDS patients, they adjusted PEEP levels randomly according to Blood Institute’s ARDS Network (ARDSnet) and SI strategy with other parameters remained consistent [113]. And SI strategy tended to set lower PEEP and had obvious lower biomarkers of lung injury (IL-6, IL-8, and TNF-a) compared to ARDSnet. The further study introduced that combining the values of plateau pressure (>25 cmH2O) and SI (>1.05) has the highest diagnostic accuracy to identify the overdistension compared to CT scan [115].

But Chiumello et al. compared PEEP levels selected by different methods in patients with different recruitability of the lung, Express, stress index, esophageal pressures, PEEP/FiO2 table [118]. Besides the PEEP/FiO2 table, other methods all provided a similar PEEP in the lung with different recruitability. But it seems unreasonable to set high PEEP for low recruitable lung.

2.6.6 The Limitations of SI

First, it just can be used in constant flow ventilation. Meanwhile, deep sedation or paralysis are required to obtain a reliable measure. Second, resistance needs to be assumed constant with inflation. Third, in situations of pleural effusions, high intraabdominal pressure, or heterogeneous lung disease, the ability of SI to identify injurious ventilation needs to be discussed.

2.6.7 Conclusion

In conclusion, a dynamic P-t curve could be used to identify injurious ventilation to some extent, but it is important to pay attention to the theoretical assumptions of applying SI.

2.7 Pressure-Volume Curves

2.7.1 Introduction

The pressure-volume curve (PV curve) is an important index reflecting the mechanical characteristics of the respiratory system. For clinicians, well understanding of the different manifestations of the PV curve under different pathophysiological conditions are helpful to guide the ventilator parameters setting in a clinical scenario and promote the clinical researches.

The PV curve first appeared in the description of acute respiratory distress syndrome (ARDS) in the 1970s [119], since then its role in diagnosis and monitoring has been gradually recognized. However, it was until 1984, when Matamis and colleagues described the relationship between the PV curve and different stages of acute lung injury in adult ARDS patients, that the curve was recognized as a potentially important tool clinically [120]. With further researches on ARDS patients’ mechanical ventilator-related lung injury and the effect of ventilator-related strategies on patients’ prognosis [121, 122], the PV curve has aroused wide interest among clinicians. By the mid-1990s, PV curves had been widely used in mechanical ventilation to reduce ventilators-associated lung injury. However, with the development of research, there was not enough evidence to prove that protective lung ventilation guided by the PV curve can significantly improve the outcome of patients. Currently, rather than directly used to guide the parameter setting of mechanical ventilation in clinical practice, the PV curve was served to enhance intensivists’ understanding of diseases and further clinical diagnosis and treatment.

In this chapter, we will elaborate the clinical applications and recent progress of the PV curve.

2.7.2 Equation of Motion

Before developing the description of applications of the PV curve in clinical practice, it was necessary to briefly understand the equation gas motion of the respiratory system. It is written as follows:

$$ \mathrm{PRS}=\mathrm{ERS}\cdotp V+\mathrm{RAW}\cdotp \dot{V}+\mathrm{PEEPTot} $$
in which PRS is the pressure in the respiratory system, ERS is the resistance in the respiratory system, V was volume, RAW is airway resistance, $$ \dot{V} $$ is airway flow, PEEPTot is the total positive end-expiratory pressure. According to this motion equation, it is no difficult to find that airway pressure of the respiratory system mainly consists of three components in patients without spontaneous breathing during mechanical ventilation. It is no doubt that airway resistance always existed during airflow movement. Therefore, accurate measurement of elastic resistance of the respiratory system, of which reciprocal also called compliance, at static state needs to minimize the influence of airway resistance. Thus, a static PV curve was also named the compliance curve. The comparison of static and dynamic PV curves will be introduced in the following paragraphs.

2.7.3 Static and Dynamic Pressure–Volume Curves

PV curves can be categorized into dynamic PV curves and static PV curves. The dynamic PV curve is a comprehensive indicator of lung/chest wall compliance and airway resistance. It is much easier to measure, but cannot reflect the overall compliance of the respiratory system accurately due to the interference of airway resistance. The static PV curve is the curve of lung volume and pressure under ideal conditions. Therefore, the static PV curve is more accurate, but also more tedious. In the actual measurement process, it is impossible to interrupt the patient’s breathing completely. Thus, it was difficult to achieve an absolute steady state of the respiratory system, and the curve we get is often referred to as the quasi-static PV curve.

At present, in the measurement of the PV curve, the Low Constant-Flow Method is widely used in clinical practice for its simplicity and easy operation, which can minimize the influence of the respiratory airway resistance on measurement results [123, 124]. At the early stage, researchers prefer the dynamic PV curve since not need to disconnect ventilators [125127]. In 2001, Adams and colleagues [128] investigated the effects of dynamic and static PV curves on different aspects of respiratory mechanics in dogs with oleic-acid-induced lung injury. They measured a quasi-static PV curve using the super-syringe method and compared it with the curves measured with the constant inspiratory flow of 10, 30, and 50 L/min. They found that as the inspiratory flow increased, the curve drifted to the right and the initial volume of all curves is positively correlated with PEEP level during the measurement of the dynamic PV curve. It was not difficult to infer from this study that when measured with a high flow, the pressure value at the same volume will leave a large pressure change, which is probably caused by airway resistance.

On the contrary, a multicenter study of 28 patients with ALI and ARDS presented a conflicting viewpoint. Stahl et al. [129] detected changes in respiratory mechanics (compliance and recruitment) during dynamic measurements with an incremental PEEP. They suggest that dynamic measurements may be more appropriate for mechanically ventilated patients than the static PV curve. Of course, the application of the dynamic PV curve is still controversial. This chapter will introduce the widely used static PV curve measurement method in detail.

2.7.4 Measurement Techniques and Mathematical models Measurement Techniques

The traditional quasi-static PV curve measurement mainly include the following three methods: super-syringe method, multiple-occlusion method, and low constant-flow method.

Super-syringe Method: The super-syringe method was first introduced in 1975. In the 1980s, it was gradually applied to the description of the different stages of ARDS [120, 121]. Before tracing the curve, the ventilator should be disconnected from the patient and a large syringe of 2–3 L and a manometer should be connected at the end of the endotracheal tube. The pressure value is recorded while pushing the syringe from FRC. Each injection of 100 mL gas is followed by a pause of 2–3 s to allow the pressure in the lungs to reach the quasi-static status to acquire the inspiratory branch of the curve. The pressure–volume value is recorded and repeated until the pressure reaches 40 cmH2O. At this point, stop the injection and gradually deflate gas in the same way, drawing the curve of the expiratory branch. Taking the pressure as the X-axis and the corresponding volume as the Y-axis, a quasi-static PV curve is drawn [130].

The virtue of this measurement method is that the quasi-static PV curves of the inspiratory phase and the expiratory phase can be recorded simultaneously. Drawbacks are as follows: (1) additional syringes are required; (2) the ventilator has to be disconnected during measurements; (3) cumbersome and time-consuming operation, which may lead to the risk of hypoxemia; (4) sedation and muscle relaxation are essential; (5) it was not easy to measure curves with different PEEPs; and (6) the continuous gas exchange, gas compressibility, and gas temperature change cannot be controlled in the measurement process. Figure 2.10 shows that the pressure-volume curve is acquired with the low constant-flow method.


Fig. 2.10

Pressure–volume curve acquired with the low constant-flow method under different levels of PEEP (inspiratory phase). The purple solid line AB means that the difference in the end-expiratory lung volume (∆EELV) between peep = 15 and zeep. The black solid line CD means that the lung recruited volume (Vrec) between peep = 15 and zeep. PEEP Positive end-expiratory pressure

Multiple-Occlusion Method: The multiple-occlusion method (also called “ventilator method”) was developed in the late 1980s and first described by Levy et al [131]. By periodically interrupting tidal breathing at different lung volumes during volume-controlled mode with a constant flow, different plateau pressures are obtained by blocking at the end of inspiratory and a quasi-static PV curve is constructed, based on the assumption that the lung relaxation volume (or FRC) remains at the same baseline during each measurement. In the early 1990s, this method was gradually applied to the determination of the low inflection point and high inflection point of the PV curve [132, 133] and the effect of different PEEP levels on pulmonary retraction in ARDS patients [134, 135]. Since this method is measured under normal ventilation conditions, it does not need to disconnect the ventilator. At the same time, the volume change caused by oxygen absorption could be negligible since each measurement lasted for 3 s. Compared with the super-syringe method, this method has less hysteresis. Meanwhile, the time consumed for each measurement is about 15 min and its accuracy is poor, which hindered the application of this method in clinical practice [125, 130].

Low Constant-Flow Method: The low constant-flow method was developed by Suratt and coworkers based on the dynamic measurement method [136, 137]. After inflate and deflate the respiratory system with a continuous low flow, a quasi-static PV curve with pressure as the X-axis and volume as the Y-axis could be drawn. Besides, capacity is calculated by integrating flow rate and time. The most suitable flow rate ranges from 3 to 9 L/min, while higher flow will cause the curve shift to the right significantly [123, 124, 135]. Compared with the super-syringe method, even the flow rate as low as 1.7 L/min might produce a certain drift [130]. Many modern ventilators equipped this function and could be performed quickly at the bedside without extra special equipment. Furthermore, this method does not need to disconnect the ventilator and no need to modify the lung volume before the measurement. However, oxygen absorption still existed in this method.

Mathematical Models: A significant interobserver variability exists in the inflection pressure point interpretation, with a maximum difference of 11 cmH2O for the same patient [139], although some study also reported good consistency [140]. To avoid these limitations, several mathematical models [141243] have been proposed. Although the fitting degree of the data models was fine, the results obtained from different data models of the same curve were also different [144].

Most researches used the model described by Venegas and colleagues [141], which was a sigmoid equation, symmetric to the inflection point. Several other models tried to improve the flexibility of the equations, but some of them have not been clinically verified. Therefore, the differences between different models should be taken into account while setting data or comparing the results of different studies based on these models.

2.7.5 Fundamental Concepts of the PV Curve

Characters of the PV Curve: The PV curve is normally measured in an upright, awake, and relaxed subject. The typical shape is generally described as alike sigmoid, consisting of three segments separated by two inflection points [140], and reflecting the balance between the lung and the chest wall. The compliance of the first and third segments is low and nonlinear. The intermediate segment between the upper and lower inflection points is considered linear and is used to calculate “linear” compliance. Changes in the linear compliance of the PV curve usually imply different stages of different diseases.

Hysteresis: The term hysteresis refers to unrecoverable energy or delayed recovery of energy which is applied to a system in physics. Since the lungs are not an ideal elastic system (in which energy changes can be recovered immediately), there is a notable difference between the inspiratory and expiratory branches of the PV curve, which increased with increasing volume and most pronounced when the volume reached FRC [145].

As early as the 1950s, researchers had recognized the surface tension as a factor of hysteresis, which verified by filling the lungs with saline and air separately on the PV curve measurement [145]. Subsequently, researchers found that small volume excursions might have more to do with the intrinsic tissue composition of the lung than alveolar surfactant [146]. Besides, hysteresis increased significantly when the alveolar collapse was caused by pulling the lungs out of the chest wall. In general, in the process of PV curve measurement, recruitment/derecruitment, surfactant, stress relaxation, and gas absorption are the main reasons leading to hysteresis [130].

Supine Posture: When the subject was in the supine position, the measured PV curve of the chest wall shifted to the right and rotated counterclockwise. FRC was reduced by half and respiratory compliance increased in the supine position compared to the supine position, possibly due to differences in abdominal compliance in different positions. Similarly, it is not difficult to understand that when patients have ascites, intraperitoneal bleeding or abdominal hypertension, etc., the abdominal compliance will be reduced, which will affect the respiratory system compliance.

Intrinsic Positive End-expiratory Pressure: Intrinsic positive end-expiratory pressure (PEEPi) is another influential factor in the measurement of the PV curve, mainly represented by the presence of positive alveolar pressure at the end of the expiratory period. PEEPi is mainly caused by shortened expiratory time or slow exhalation due to high airway resistance, resulting in higher alveolar pressure at the end of an exhalation than the endotracheal pressure and incorrect measurement of respiratory compliance. In other words, when PEEPi is present, the pressure difference between the end of exhalation and the end of inhalation is overestimated, leading to an underestimation of respiratory compliance [148]. Besides, PEEPi has been shown to contribute to a portion of re-expanded alveolar hyperventilation. Therefore, it is important to empty the lungs by extending the exhalation time before the PV curve is measured.

Lung Versus Chest Wall: The PV curve of the respiratory system is composed of two main parts: the lung and the chest wall. The chest wall compliance is equal to the change in volume divided by the change in trans-chest pressure measured by the esophageal pressure tube, and the lung compliance is equal to the change in volume divided by the change in transpulmonary pressure. As compliance changes, it is important to recognize the extent to which both are present in different diseases. The relationship between the three is as follows:

Jul 31, 2021 | Posted by in RESPIRATORY | Comments Off on Mechanics
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