As a guide to airway management, this book has so far focused mainly on establishing a secure channel between the trachea or main bronchi and the outside world, usually in the form of an endotracheal tube. This is a worthy goal and a prerequisite for life for most people, but it remains a means to an end. After the tube is in place, we must do something useful with it, namely, ventilate and oxygenate the patient. This chapter will introduce readers to positive pressure ventilation (PPV) by discussing basic concepts and definitions in respiratory physiology and universal themes in ventilator design. As we look at each of several classic modes of ventilation, we will suggest typical applications, guidelines for initial setup, and discuss the peculiarities of each mode. Finally we will briefly summarize some of the more recently developed approaches to mechanical ventilation in difficult to ventilate or oxygenate patients.

Ventilation is simply mass movement of gas in and out of the lungs, or breathing. Regardless of whether an individual is breathing spontaneously or with artificial PPV, the volumes of gas moving in and out of the lungs can be measured in a process called spirometry. When these measured volumes are added up, we refer to the resulting sums as lung capacities.

Figure 54-1 shows a graph of volume versus time for respiration, which begins as normal relaxed breathing or tidal breathing. The volume moving in and out with each normal breath is called tidal volume (V_T). At time “a” on this graph, the person has inspired (breathed in) as much as possible and then at time “b” he has gone on to expire (breathe out) as much as possible. Inspiratory reserve volume (IRV) and expiratory reserve volume (ERV) are the maximum inspiratory and expiratory volumes, respectively, that can be attained beyond tidal ventilation. Residual volume (RV) is the volume remaining in the lungs that cannot be forcibly expelled during active expiration. The total volume that remains in the lungs after relaxed tidal expiration is the sum of ERV and RV and is called functional residual capacity (FRC). The total volume that can be inspired from that point is the sum of V_T and IRV and is called inspiratory capacity (IC). The total volume that a person can move in and out of the lungs at maximum effort is the sum of ERV, V_T, and IRV and is called vital capacity (VC). The total volume within the lungs at maximum inspiration is the sum of all the lung volumes, and we call it total lung capacity (TLC).

Like all moving fluids, the gases that we breathe are pushed from a region of high pressure toward a region of lower pressure. In the case of spontaneous breathing, the diaphragm and other respiratory muscles work together to expand the thorax and create negative intrathoracic pressure relative to the outside atmosphere (note that throughout this chapter and in most of physiology, pressure will be measured relative to atmospheric pressure and atmospheric pressure is considered to be zero). In the case of most artificial ventilation, we apply positive pressure at the upper airway or within the trachea, and gas moves toward the region of relatively lower pressure within the lungs. In both situations, expiration is mainly passive. It is driven by lung elasticity pulling the intrathoracic volume down and increasing intrathoracic pressure until, at the point of full exhalation during tidal breathing (FRC), the inward force of lung elasticity is equal in magnitude and opposite in direction to the outward recoil of the relaxed inspiratory muscles and other tissues of the chest wall. Additional active expiration may be accomplished by contraction of the internal intercostal muscles and the muscles of the abdominal wall.

Before we delve into the specific topic of mechanical PPV, it is worth mentioning that there are other ways to provide artificial ventilation. First, there is the largely
historical example of negative pressure mechanical ventilation such as that provided by the iron lung. This device surrounded the patient’s thorax while a seal excluded the neck and head. As the bellows, or, later, a separate compressor, generated negative pressure in the tube around the thorax, the chest wall transmitted it to the lungs and air flowed into that low-pressure region, generally by way of the natural airway. It served many patients with neuromuscular respiratory failure for decades of their lives, but it is difficult to move about to say the least, and even difficult to provide nursing care when the patient’s body must be encased by the ventilator, so it has now largely been replaced by PPV, even for patients with very long-term ventilator dependence.

The other frequent exception to the model of mechanical PPV is that of manual PPV as provided by anesthesia and critical care practitioners daily by means of a gas-filled bag attached to a one-way valve. Providing manual breaths via natural or artificial airway is a lifesaving skill and can lead to a more intuitive understanding of the principles of PPV. Nonetheless, even the most seasoned anesthetist will fatigue at some point, so most of us will enjoy the luxury of being able to replace the work of our hands and forearms with a machine that is expressly made to ventilate lungs. This brings us to the remainder of the chapter, which will explain principles and practice of mechanical PPV.

Any reader who is not a respiratory therapist will likely recall his or her first introduction to mechanical ventilation as an alphabet soup of acronyms and confusing definitions. Unfortunately, naming and availability of modes of ventilation is inconsistent between different ventilator manufacturers. Furthermore, with the advent of computer-controlled ventilators, the subtle variations in modes of ventilation have become potentially limitless. For this reason, it is worthwhile to pause and build a strong theoretical framework for how we describe ventilator function and control. Unlike marketing pitches, the principles that govern the flow of gas in and out of human lungs are fairly simple and immutable. Fundamentally, any mechanical ventilator is a machine that uses a pneumatic or electric power source to take over a patient’s work of breathing. Naturally, this must occur in a carefully controlled fashion in which some relevant variables related to respiration can be measured and input into the ventilator control system, which will process them and modify ventilator output accordingly. If the reader can understand all the possible ways that someone might design a machine to interpret and interact with these laws of fluids and respiratory mechanics, it should be easy to adapt to small modifications in the currently existing technology as they come about.

Regarding the power source, most modern ventilators use some combination of electrical and pneumatic power. Since stored medical gas in hospitals is already pressurized to approximately 50 psi, and gas stored in smaller tanks is at much higher pressure, it is efficient to use some of the energy that is released in the process of expanding stored gas to drive a bellows. Alternatively, some devices simply use regulator valves that only decrease the pressure down as far as the desired airway pressure mandates. On the other hand, the power source for a transport ventilator or the backup power source for a stationary ventilator is typically an electric compressor. Similarly, the control system for a simple ventilator can be purely mechanical and pneumatic as is the case for many of the simple intermittent positive pressure breathing devices that are used to administer inhaled medications. Modern ventilators, however, almost exclusively use electricity for their control systems: alternating current circuits for the small motors and direct current circuits for the computer systems and sensors.

The desired output of a mechanical ventilator is obviously ventilation. More specifically, appropriate ventilation must provide adequate minute ventilation.
Appropriate ventilation can minimize both dead space ventilation (the amount of ventilation that does not participate in gas exchange because it is only in poorly perfused parts of the lung or airway) and shunt (blood that cannot participate in gas exchange because it flows through parts of the pulmonary circulation that are not well ventilated). For any patient there is an ideal balance point. We must provide large enough V_T so that all perfused portions of lung get ventilated and V_T is much larger than the wasted anatomic dead space ventilation, but we must avoid applying so much pressure to the lungs that it impedes pulmonary blood flow and increases physiologic dead space. In addition, excessive positive airway pressure can be harmful to the tissue of the lungs and reduce cardiac output, so the goals of ventilation must be balanced against the desire to limit peak airway pressure. The ideal rate of gas flow into or out of the lungs can also vary between individuals based on traits such as level of sedation, respiratory drive, body habitus, or specific lung or airway disease states.

The three main variables of PPV, namely volume, pressure, and flow, can be related to one another using a simplified model of respiratory mechanics. In this model, we will imagine all of the airways, bronchi, and bronchioles being represented as a single tube leading to a balloon that represents all of the combined alveoli (Fig. 54-2).

Using this single alveolus model, we can see that during a positive pressure breath, the inspired gas would be driven into the alveolus from a region of high pressure at the proximal airways toward a region of lower pressure in the alveolus, but we can also imagine that this movement would be opposed by the dynamic resistance to the flow of gas through the tube. The relationship between pressure, flow, and resistance is described by the following equations.

This relationship is dynamic because it involves a rate, and it only applies when there is gas flow through the tube. If there is no flow, then resistance is irrelevant, and the pressure that we measure at the airway cannot be influenced by it. Looking back at our balloon-on-a-straw however, we can also easily understand that there is a second, static force that opposes filling the balloon, and that is the elastic properties of the expanding balloon itself. The pressure that the balloon will exert back on the gas within it is governed by its elastance and volume according to the following equations:

If pressure at the mouth of the tube is set at a given amount and ample inspiratory time is allowed, elastance will determine what volume enters the balloon. When the volume is reached at which back pressure from the elastance of the balloon equals airway pressure, flow will stop. Then if the positive airway pressure is reduced or removed, the direction of the flow will be reversed until the pressure in the balloon is again equal to airway or ambient pressure. Note that we often discuss elasticity in terms of its inverse, compliance, which is simply change in volume over change in pressure.

Combining the static and dynamic components of the work of breathing additively, we arrive at a simplified equation of motion for the respiratory system.

This equation of motion allows us to relate all of the key variables of respiratory mechanics. Each of the variables can be manipulated in one way or another. Resistance and elastance are generally characteristics of the patient, whereas pressure, flow, and volume are each variables that that may be set by the ventilator.

Resistance to flow of an ideal fluid is described by Poiseuille’s Law.

Resistance = (8 × viscosity × length) / (pi × radius⁴)

The natural airways are difficult or impossible to describe this way because they taper and branch and the flow in most of the airways is turbulent, but Poiseuille’s Law is still a worthwhile model to keep in mind. It forms the theoretical basis for using a shorter, larger radius endotracheal tube or giving bronchodilator medications to reduce the resistance to airflow, or even the practice of mixing helium into the inspired gases for patients with severe airway obstruction in the hope of reducing resistance by reducing fluid viscosity of the inspired gas itself.

Like resistance, elastance in a living human is more complex than our single alveolus model would suggest. Rather than a single balloon, the intrinsic lung elastance is a function of several hundred million alveoli and the tissue that forms them, as well as the airways themselves. The elastance of real lungs is not constant either. Rather, it increases with increasing lung volume (Fig. 54-3).

In healthy people with normal lung tissue, a more significant component of elastance is often related to extrinsic sources. These include the pressure of an obese or insufflated abdomen pressing up against the diaphragm, variations in patient position such as Trendelenburg or prone positioning, or forces from the tissues of the walls of the thorax including active movements of the respiratory muscles, which may aid or oppose the ventilator. These factors are all important to consider, and they can be modified by a human clinician, but they are essentially constants from the perspective of the ventilator control algorithm. In contrast, the flow, volume, and pressure components of the equation of motion form the heart of ventilator management.

As mentioned above, there are a plethora of specific modes of mechanical ventilation, and we will only discuss a few of the classics here. Before we proceed with this discussion, though, it will be valuable to consider the ways a ventilator could be controlled in theory, because most of these approaches are indeed being used by one device or another.

First, what can a ventilator know? That is to say, what variables are commonly input into a control algorithm? For a start, a ventilator must monitor the three essential parameters that will determine the mass movement of gas as discussed above: pressure, flow, and volume. Like so many other things, the exact location of these sensors (eg, inspiratory arm, expiratory arm, T-piece next to endotracheal tube itself) and subsequent way that the data is used varies between manufacturers. As a curiosity, it is also notable that flow is the derivative of volume with respect to time, and, inversely, volume is the integral (or area under the curve) of flow over a period of time, so many ventilators measure one of these two variables and calculate the other. If the ventilator does not have a bellows or piston, most likely it is actually measuring flow and integrating it to produce the values of volume that show up on its settings or display screen. We will intentionally perseverate on pressure, flow, and volume, but perhaps the most fundamental variable that the ventilator keeps track of is time. Almost any other variable for which an electronic sensor has been invented is fair game for the ventilator to monitor and respond to, and there are a few notable examples. Many modern anesthesia ventilators integrate the concentrations of several inspired and expired gases into their displays, and even many intensive care ventilators measure and display or trigger alarms based on inspired O₂ and expired CO₂ levels. It is also possible to use pressure sensors outside of the ventilator circuit or even electrodes on the chest that measure impedance to estimate chest volume. Indeed, both of these techniques are used clinically to trigger the inspiratory phase in some infant ventilators. Of course, ventilators must also be able to act on human operator inputs and most of their functions can be triggered manually by a clinician. As any of these variables are measured and input into the ventilator control system, they can be used for different purposes, and we will categorize them further based on what the control algorithm does with them.

A control variable is simply the parameter that the ventilator controls during inspiration. Its magnitude could
be constant, but just as often it varies over the inspiratory time as is demonstrated by the many waveforms in Fig. 54-4. Note, for instance, that the first pressure control waveform applies constant pressure, but the second uses an ascending ramp waveform for pressure. Both of these examples are still using pressure as the control variable though. In mathematical terms, the control variable is the independent variable in the equation of motion, and at any given time, the other two ventilator-specific variables depend on the interplay of the control variable with the patient-specific properties of elastance and resistance. For example, if we set a particular pressure waveform as the control variable, then VT and flow rates will result based on the resistance and elastance of the system. They are directly related to each other, and if the pressure is increased, higher volume and flow will result, whereas if the pressure is set lower, there will be a decrease in the volume and flow delivered. On the other hand, if a mode of ventilation that specifies volume or flow for the control variable is chosen, then we cannot have direct control over the pressure that is generated to reach that target, but we still know that the variables will be directly proportional to one another. As noted previously, volume and flow are closely related, with flow being the derivative of volume with respect to time, so for clinical purposes when we control one of these two variables over a set period of time, we are, in essence, controlling the other.

With regard to time, our simplified equation of motion explains some important relationships, but it does not tell us much about how they change over the time of a respiratory cycle. First, recall that elastance increases with increased lung volume, so during inspiration, elastance increases with time also. In contrast, resistance is usually higher at lower lung volumes, but the magnitude of that change is smaller over a typical respiratory cycle, so we can usually ignore it. The remaining relationships will always be consistent with the equation of motion, but we can set various waveforms for the control variable. Some common control waveforms and typical resulting waveforms of the dependent variables are shown in Fig. 54-4.