Evaluation of Pulmonary Function in the Neonate

Introduction

The ability to measure lung function is essential for understanding lung growth and developmental physiology, diagnosing pathologic conditions, and assessing therapeutic interventions and management. Neonatal pulmonary function evaluation in the 21st century has achieved a remarkable level of sophistication, making use of new and emerging technologic innovations; these advances stand in stark contrast to very early lung function evaluation efforts that occurred in the 19th century with the invention of a simple spirometer to measure vital capacity. ^, Modern pulmonary function testing in adults and children has been used effectively as a research and clinical tool since the 1950s. ^, Although testing of infant pulmonary function was introduced around the same time, its practical use lagged behind, owing to technical limitations, lack of patient cooperation, difficulty with the long duration of testing in this age group, and time-consuming manual analysis methods. In the 1980s, improvements in technology, including sensor miniaturization and computerization, led to mobile equipment systems enabling the clinician to perform routine pulmonary function testing at the bedside. In the 1990s, a new generation of infant ventilators incorporating continuous display of pulmonary function data, as well as online graphic display of respiratory waveforms, became available. Infant pulmonary function testing has thus moved from the research laboratory to the patient bedside, while efforts to reduce invasiveness and improve accuracy continue. This chapter will give an overview of the basic principles and methods for evaluation of pulmonary function in the neonate from both a physiologic and technical perspective.

Physiologic Background

Each respiratory cycle is governed by a driving pressure that moves a volume of air into and out of the respiratory tract, resulting in measurable airflow and volume changes. During unaided breathing, the driving force required for initiation of inspiratory airflow is generated by contraction of the respiratory muscles, causing a downward movement of the diaphragm and outward movement of the rib cage, resulting in a transient decrease in alveolar pressure from atmospheric pressure at end-expiration to a sub-atmospheric level and peak airflow at mid-inspiration ( Fig. 68.1 ). During expiration, relaxation of the respiratory muscles and inward recoil of the lungs result in an increase in alveolar pressure, causing an expiratory airflow that reaches a peak near mid-expiration. At the end of expiration, alveolar pressure returns to atmospheric level, airflow returns to zero, and another respiratory cycle commences. Intrapleural pressure, however, remains sub-atmospheric throughout the breathing cycle owing to a balance of the opposing forces of lung elastic recoil and chest wall outward recoil.

Fig. 68.1, Pressure, flow, and volume signals for a respiratory cycle during spontaneous breathing.

During tidal spontaneous breathing in normal healthy lungs, pleural pressure decreases and increases with inspiration and expiration, respectively. The magnitude of these changes is determined by airflow and tissue resistances and by elastic forces in the lungs and chest wall, reaching a maximum at inspiratory and expiratory midpoints, as do the inspiratory and expiratory airflows.

Evaluation of pulmonary function is based on measurements of the interaction of driving pressure ( P ), volume ( V ), and airflow ( $\dot{V}$ $\dot{V}$ ). Early investigators developed an equation of motion relating the movement of gas, chest wall, and lungs in a force balance equating the total driving force for breathing to the sum of elastic, resistive, and inertial forces ^, :

P = E V + R \dot{V} + I \ddot{V}

$P = E V + R \dot{V} + I \ddot{V}$

where E = elastance of the respiratory tract, or its tendency to recoil upon application of an expanding or compressing force; equivalent to the reciprocal of compliance, or 1/ C , R = frictional resistance of respiratory tract comprising both airflow and tissue resistances, and I = inertance of the respiratory tract, due to gas and tissue acceleration, usually considered negligible except at higher breathing frequencies and flow rates, V = volume, $\dot{V}$ $\dot{V}$ = flow (rate of change in volume, $d V / d t$ $d V / d t$ ), and $\ddot{V}$ $\ddot{V}$ = acceleration (rate of change in flow = $d^{2} V / d t^{2}$ $d^{2} V / d t^{2}$ ). ^,

This equation describes a one-compartment linear model relating driving pressure to volume and flow and has been considered an ideal elemental equation as well as a basic principle of respiratory physiology. However, it is well appreciated that the respiratory system is not a linear system and that resistance, compliance, and inertance are not constants, being dependent on lung volume, volume history, flow characteristics, and breathing frequency. The respiratory system is a complex system consisting of multiple components with differing mechanical properties. Despite its inexactness, however, and because of the complexities of dealing with nonlinear modeling, this simple linear model provides a useful framework for examining the dynamic behavior of the normal respiratory system.

Instrumentation

Instrumentation to measure the pressures, gas flows, and volumes in neonatal pulmonary function testing must meet basic physical capabilities for accuracy, precision, range, frequency response, and calibration integrity. ^, Computerized data collection systems also need to conform to signal processing standards, such as sufficient digitization resolution and adequate sampling rates to minimize aliasing in digitized signals and to permit the use of appropriate automated breath detection algorithms. ^, In addition, devices to be used in line with infant breathing circuitry also must meet minimum standards for safety, hygiene, dead space, and resistive load. Equipment specifications and methodologies have been studied extensively in an effort to achieve uniformity among research laboratories and also among the various commercially available infant pulmonary function assessment systems. To this end, comprehensive testing and measurement standards have been established and published. ^,

Flow Measurement

Decades of pulmonary function research have advanced several flow measurement technologies for use in both term and preterm neonates. In spontaneously breathing, nonintubated infants, airflow is measured using a flow sensor attached to a face mask, whereas in intubated infants, the flow sensor is placed in line with the endotracheal tube at the ventilator connection. A consistent objective has been to produce a smaller and lighter flow sensor having minimal dead space and airflow resistance. Some of the most commonly used flow measurement devices are described below.

Pneumotachometers

Pneumotachometers are flow-resistive type devices in which gas flows through a tube containing a fixed laminar flow-resistive element. The resistive element can be either a fine-mesh screen or a bundle of small capillaries (Fleish type), both of which produce a pressure drop that is linearly proportional to flow so long as it is within a specified laminar flow range; higher flows give rise to turbulence and a nonlinear response. ^, Accordingly, pneumotachometers that have linear flow ranges appropriate to the maximum expected flow for a given subject should be selected. Condensation of water vapor on the resistive element can easily alter its resistive properties, so most pneumotachometers incorporate a heating element to prevent accumulation of moisture. The flow characteristics may also be altered by accumulation of secretions, varying gas viscosity, gas composition, and temperature. ^, Pressure drop across the resistive element is measured with a sensitive differential pressure transducer, which must have a linear response over the appropriate pressure range and with sufficient frequency response and phase characteristics to capture any rapid transients contained in the flow signal. ^, The pneumotachometer, as well as most other types of flow sensors, should be calibrated using the same connectors, adapters, and gas composition as will be used during actual measurements at the bedside.

Hot Wire Anemometers

Hot wire anemometers are flow sensors that operate on the principle that the electrical resistance of a metal is temperature dependent. In one application of this principle, the amount of electrical current needed to maintain a constant temperature in a fine heated wire suspended across an air stream is measured and related to the magnitude of the airflow; the added current increases as the airflow increases and more heat is dissipated. The response of the anemometer is inherently nonlinear, requiring linearization either in the signal conditioning circuitry or through computer signal processing. Another disadvantage is directional insensitivity, usually circumvented by the use of multiple sensor elements. Advantages are small size, good frequency response, and wide dynamic range. Furthermore, hot wire anemometers are relatively pressure-independent devices and thus are not subject to artifacts from sudden pressure transients, as are pressure-dependent devices. ^,

Other Flow-Resistive Type Flow Sensors

These include the fixed and variable orifice types and various pitot tube configurations. These flow sensors also require a pressure measurement, but, unlike the laminar flow pneumotachometer, the flow-pressure relationship for these devices is nonlinear and requires hardware or software linearization. Like the anemometer, these sensors typically are smaller and lighter and thus have less dead space than the pneumotachometer. These factors, in addition to being less susceptible to moisture condensation errors, allow these devices to be more appropriate for long-term monitoring, especially as integral components of mechanical ventilators for flow control and tidal volume monitoring, as well as for breath detection and triggering.

Ultrasonic-Beam Flow Sensors

The theory behind operation of the ultrasonic-beam flow sensor is based on the fact that sound waves traveling through a medium in flux are either accelerated or slowed in accordance with the velocity of the medium. By measuring the transit time of a sound wave pulsed through a gas stream in a rigid conduit, the flow rate of the gas can be calculated. The main advantage of the ultrasonic-beam flow meter is a much lower sensitivity to changes in gas properties than is associated with a traditional pneumotachometer, but a major disadvantage of this flow meter is a larger dead-space volume. ^,

Emerging Technologies

Two notable new classes of flow sensors are electro-optical and micro-machined sensors, both with various implementations of their basic operating principles. One variation of an electro-optical–type sensor is based on measurement of the deflection of a light beam emitted from a short length of optical fiber suspended in an air stream. Micro-electro-mechanical systems–type sensors integrate micro-machining technology with micro-electronics and are based on measurements of mechanical deflection, thermal effects, or a combination of both. An example of a mechanical-type sensor utilizes a micro-cantilever machined on a silicon chip. Bending or vibration of the cantilever in a flow stream causes a change in a piezo-resistive element, which is transduced into an electrical output. In addition to the low cost of manufacturing, advantages of these newer sensors are small size, low dead-space volume, fast response, wide dynamic range, and negligible resistive loading.

Volume Measurement as Derived From a Flow Signal

When respiratory airflow is measured directly at the airway opening using a flow sensor, volume change in the lungs may be derived by integration of the flow signal over time, which is simply the area under the scalar flow waveform. With computerized signal processing, electronic integration of the flow signal has been largely replaced by digital integration. Sign reversals in the flow signal are used to detect inspiratory and expiratory endpoints, which are then used to identify individual breaths. Typically, the resultant tidal volume waveform will have slightly unequal inspiratory and expiratory portions owing to differences in inspired versus expired conditions in gas composition, water vapor, temperature, and viscosity. For these and other reasons, most volume signal conditioning routines correct for any baseline drift by incorporating an automatic return to zero at end-expiration. Differences in inspiratory and expiratory tidal volumes larger than 10%, however, may be indicative of airflow leakage around the endotracheal tube or face mask. Large baseline drifts may also indicate faulty flow sensor readings, which may be due to factors such as water vapor condensation or buildup of secretions.

Indirect Volume Measurements

Indirect methods for measuring volume change, without connection to the airway opening, have also been implemented. These are useful, for example, when an infant is receiving non-invasive forms of respiratory support such as nasal continuous positive airway pressure (nCPAP) or heated and humidified high flow nasal cannula (HFNC). In these instances, the requirement of a nasal interface to apply the therapy precludes use of a traditional direct airflow sensor connected to a face mask. As a workaround, alternative techniques using measurements at the body surface instead of at the airway opening may be used to derive not only tidal volume but also various other indices of respiratory function. Advantages include elimination of face mask and flow sensor dead space and airflow resistance, which affect the actual parameters being measured. These indirect measurement devices also have the advantage of allowing noninvasive continuous monitoring without alterations in the breathing pattern that occur with use of face masks or inline flow sensors, and are therefore especially useful for undisturbed measurements in spontaneously breathing and sleeping infants. ^, Some of these technologies are described below.

Respiratory Inductance Plethysmography

Respiratory inductance plethysmography (RIP) was first developed in the early 1980s and has been commonly used for chest wall motion analysis and for monitoring breathing in sleep apnea studies, as well as for non-invasive spirometry. The basic RIP apparatus consists of sinusoidally arranged thin wires embedded in two soft elastic bands, which are placed around the rib cage and abdomen as shown in Fig. 68.2 . As a patient breathes and the bands expand and retract, changes in the magnetic field created by a low-level high-frequency current in the coils cause changes in their self-inductance. It follows that these inductance changes are proportional to changes in the cross-sectional areas circumscribed by the bands. The resulting small changes in frequency are transduced into electrical signals, which may then be recorded. Rib cage (RC) and abdominal (ABD) excursions are measured separately and the summation of these two signals (SUM) is approximately proportional to the net pulmonary tidal volume change.

Calibration of the SUM signal to true tidal volume can be accomplished using one of several techniques that have been developed for this purpose by various investigators. ^, In general, all such calibration schemes require a period of restful breathing to generate separate RC and ABD weighting factors to account for their different contributions to the SUM. A period of breathing through a calibrated face mask-attached flow sensor, such as a pneumotachometer, is then used to convert the final corrected SUM signal to known volume changes and that can then be converted to flow through mathematic differentiation. Once calibrated, maintaining consistent accuracy is highly dependent on maintaining both body position and prevention of band slippage.

Nonetheless, the RIP technique, when accurately calibrated and used in place of a mask-attached flow sensor, has been used to measure undisturbed tidal volume in respiratory assessment studies for bronchopulmonary dysplasia in VLBW premature infants and in infants receiving non-invasive respiratory support. Moreover, calibrated RIP has also been shown to yield clinically acceptable pulmonary mechanics measurements in infants on nCPAP or HFNC. RIP technology has gone through several iterations in the last decades, the most recent being a wireless system (PneuRIP) that displays and records real-time measurements of several work of breathing and breathing efficiency indices (as described later in this chapter) on a small hand-held display. ^,

Electromagnetic Inductance Plethysmography

Electromagnetic inductance plethysmography (EIP) is a novel technology that measures changes in an electromagnetic field induced in coils carrying a weak high-frequency current and embedded in an elastic vest encircling the infant’s thorax and abdomen. An antenna positioned above the infant detects magnetic field changes that are proportional to cross-sectional respiratory volume changes from the chest and abdomen. ^, Advantages are subject-independent calibration and continuous undisturbed display of separate rib cage and abdominal volume excursions without the need for a face mask and flow sensor. Accuracy is dependent on proper fit of the elastic vest and the requirement for the patient to remain in a supine position with minimal hip flexion. In addition, the EIP system must be positioned to avoid nearby electromagnetic interference. EIP has been used successfully for infant spirometry as well as for measuring tidal breathing parameters with adequate clinical accuracy in term and preterm infants.

Optoelectronic Plethysmography

Optoelectronic plethysmography (OEP) is a technically sophisticated technology that requires powerful computer software to process multi-camera video images from a grid-array of reflective markers strategically placed on an infant’s thoracoabdominal surface. Markers are placed on both anterior and lateral surfaces. The number of markers can vary from 24 for neonates up to 89 markers for adults; accuracy increases as a function of the number of markers used. Respiratory volume is estimated by measuring the three-dimensional position of each marker using 4, 6, or 8 special video cameras and processing the signals through a motion analyzer. ^, The system measures volumes for three thoracoabdominal compartments, where the rib cage compartment is partitioned into pulmonary rib cage (RC _p ) and abdominal (RC _abd ) volumes in addition to the abdominal volume (ABD). These are further partitioned into left- and right-side measurements, or hemithoraxes. Synchrony and timing parameters between the various compartments can then be further analyzed separately. Studies with term and preterm infants have demonstrated accurate measurements of tidal volume as validated against pneumotachometry and showing good agreement.

Advantages of the OEP technique include its non-invasiveness and the ability for continuous uninterrupted monitoring and not requiring pre-calibration. In contrast to other indirect technologies, OEP is able to estimate not only tidal volume but also the total absolute thoracoabdominal as well as the various separate component volumes. The main disadvantages are the complex and cumbersome bedside instrumentation and the need for a large number of markers on an exposed skin surface, which also need to be precisely positioned in order to ensure accuracy. Nonetheless, OEP is a promising new technology for both research and clinical application enabling sophisticated and highly detailed measurements of a three-compartment model and their interaction.

Structured Light Plethysmography

Structured light plethysmography (SLP) is a more recently developed three-dimensional optical, non-contact, technology. As with OEP, computer vision and processing is used to generate a three-dimensional reconstruction of the thoracoabdominal surfaces. But instead of tracking markers placed on the body surface, SLP uses a checkerboard grid light pattern that is projected onto the patient’s frontal thoracoabdominal surface. As the patient breathes, movement of distinct points in the light pattern are tracked by two cameras and image processing algorithms are used to display a three dimensional spatial image of the thoracoabdominal surface displacement. The system can differentiate rib cage and abdominal volume changes as well as left and right side hemithoraxes over time. The current state of SLP technology is not yet configured to output absolute volume measurements and thus the system does not require calibration prior to use. As with most non-invasive methodologies, accuracy is motion- and position-sensitive, requiring some patient cooperation. Therefore SLP has been used chiefly for measurements of tidal breathing timing indices in children and adults, but recent studies using SLP for measurements on neonates have also demonstrated promising results. Integration of the light projector, cameras, and computer into a stand-alone unit has allowed a commercially available version of this device to be suitable for cordless bedside measurements.

Electrical Impedance Tomography

Electrical impedance tomography (EIT) is an advanced imaging technology that measures the bioelectrical impedance within the thorax. Electrical conductivity varies for different materials such as air, blood, and tissues, with air having much less conductivity than the surrounding fluids and tissues. It follows that measurement of the impedance in a transthoracic cross-sectional slice using a series of surface electrodes can provide a two-dimensional image of the impedance distribution in the cross-section. Implementing this principle, EIT is used to image the pulmonary ventilation distribution over time by measuring impedance across sequentially alternating pairs of electrodes placed circumferentially across a given transthoracic cross-section. ^, This process may be performed both statically and dynamically to monitor the breath-by-breath ventilation distribution and regional lung perfusion as a function of time. An electrode belt of as many as 16 electrodes must be placed completely around the thorax for these measurements.

Eectrical Impedance Segmentography

A closely related technology developed into a commercially available device is electrical impedance segmentography (EIS). EIS measures impedance changes in the four thoracic quadrants encompassing the upper left, upper right, lower left, and lower right pulmonary regions ( Fig. 68.3 ). Four ventral and four dorsal current collection electrodes are placed in opposition across each of the four thoracic quadrants and a fifth ventral and dorsal current injection electrode is placed centrally along the sternum. A real-time computer display of the relative percent ventilation distribution in each quadrant is provided along with a scalar display of the tidal impedance changes for each breath. An animal study has demonstrated accurate calibration of tidal impedance to known pneumotachometer tidal volumes in milliliters, but this ability has not been repeated in infants. EIT and EIS have demonstrated usefulness in evaluating efficacy of respiratory support strategies, detection of pneumothoraxes, pulmonary recruitment maneuvers, and the study of ventilation inhomogeneity. Limitations of this technology include low spatial resolution and effects of electrode placement difficulty on accuracy.

Fig. 68.3, Electrical impedance segmentography (EIS) system showing ventral and dorsal electrode placement and a simplified computer display showing breath-by-breath impedance changes as both scalar traces and as percentage ventilation relative to a baseline starting value in the four pulmonary quadrants of the thorax. (Angelie EIS system, EMS Biomedical, Korneuburg, Austria.)

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