ESTIMATION OF THE MECHANICAL PROPERTIES OF THE …

ESTIMATION OF THE MECHANICAL

PROPERTIES OF THE AIRWAYS AND

RESPIRATORY TISSUES IN INFANTS

BY LOW-FREQUENCY FORCED

OSCILLATIONS

Graham Langley Hall, BAppSci.

This thesis is presented for the degree of Doctor of Philosophy of the

University of Western Australia

Department of Paediatrics

2000

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ABSTRACT: Infant lung function tests commonly assess forced expiratory flows and/or volumes. These

techniques are unable to provide information on the site of changes in lung mechanics.

Preliminary studies using the low-frequency forced oscillation technique (FOT) have shown

the technique able to determine the low-frequency impedance spectra (Zrs). A model,

containing frequency independent airway and constant-phase tissue compartments can be

fitted to Zrs and provide an assessment of both the airway and respiratory tissue mechanics.

Aims: The aims were: (1) use the low-frequency F O T to partition Zrs into lung (Zl) and

chest wall (Zw) impedances and subsequently into nasal (Zn) and lower respiratory system

impedances (Zlrs); (2) characterise the relationship between the airway and respiratory

tissue mechanics and length in normal infants; (3) compare the airway and respiratory tissue

mechanics in wheezy infants with those of normal infants; and (4) assess the methacholine

responsiveness in infants using the low-frequency FOT. Results: Longitudinal

measurements of the airway and lung tissue mechanics over a period of 14 days were

carried out in a Brown Norway (BN) rat model allowing animals to act as their o w n

controls. A n esophageal catheter was used to non-invasively partition Zrs into Zl and Z w

and the airway and tissue mechanics of Zl and Z w described. In five mechanically ventilated

patients Zrs was non-invasively partitioned into Zl and Zw. The chest wall was found to

contribute a negligible amount to the respiratory system resistance (Rrs) and inertance (Irs),

but to contribute a significant amount to the respiratory tissue damping (Grs) and elastance

(Hrs) (38.5±7.3% and 34.4±7.4%; mean±SEM, respectively). In spontaneously breathing

infants Zn influenced only the airway properties of the respiratory system, contributing

44.6±4.9% of the Rrs and a majority of the Irs (71.7±3.5%). Forced expiratory volumes

(FEV0.5) and low-frequency F O T mechanical parameters were determined in a cross-section

study of healthy infants and the relationship between the lung function parameters and

length investigated. Both airway and tissue parameters demonstrated a decreasing quadratic

relationship with increasing length, while FEV0.5 showed an increasing cubic relationship

with length. A family history of asthma was found to have a negative effect on airway

resistance (Raw), Hrs and FEV0.5, whereas gender did not have any effect. A similar study

carried out in 22 asymptomatic infants with a history of recurrent or persistent wheeze and

standardised variants (Z scores) were used to compare the wheezy population to the

previously described normal population. Abnormal mechanics of the airways and respiratory

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tissues were found in infants with recurrent wheeze persisting beyond 1 year of age,

suggesting that infantile wheeze is not merely related to airway function, but rather an

alteration in mechanical characteristics of both the airways and the respiratory tissues. The

responsiveness to inhaled methacholine (Mch) in 17 infants was assessed, by comparing the

concentration of M c h causing an increase in Raw, equivalent to twice that of the baseline

variability (TCRaw) with the concentration of M c h causing a 1 5 % decrease in FEV0.5

(PC15FEV0.5). T C R a w was significantly lower than PC15FEV0.5, with the difference being

attributed to the relative changes required to reach the respective responses. A n increase

in R a w of 3 0 % was found to equate to the same relative change required to cause a

decrease in FEV0.5 of 15%. The concentration causing an increase in R a w of 3 0 %

(PCsoRaw) and PC15FEV0.5 were subsequently found to be equal (1.5; 0.5-2 m g / m L and

1.5; 1-4 mg/mL (median;25-75% CI), respectively). Significant changes in Irs, G and

hysteresivity were also recorded. Conclusions: The low-frequency F O T is able to assess

both the airway and respiratory tissue mechanics in sedated infants. The mechanical

properties of the airways and respiratory tissues in infants with no respiratory disease were

characterised, both airway and respiratory tissue mechanics were significantly abnormal in

infants with recurrent wheeze. The technique was found to be useful in assessing the

responsiveness to inhaled M c h in infants. Future studies should consider the use of either

a nasal or esophageal catheter to detect changes in airway or parenchymal mechanics,

respectively.

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TABLE OF CONTENTS: ABSTRACT: 3

ACKNOWLEDGEMENTS: 11

ABBREVIATIONS: 13

NOTATIONS: 15

STATEMENT OF CONTRIBUTION: 17

CHAPTER 1: REVTEW OF THE CURRENT LITERATURE 19

1.1 Overview of respiratory physiology 19

1.1.1 Basic function and mechanics of the airways 19

1.1.2 Basic function and mechanics of the lung parenchyma 20

1.1.3 Modelling the respiratory system 21

1.1.3.1 Resistance 22

1.1.3.2 Elastance ..- 23

1.1.3.3 Inertance 24

1.1.3.4 Models of the respiratory system 24

1.2 Measurement of respiratory mechanics in infants 31

1.2.1 Measures of forced expiratory flow 32

1.2.2 Measures of compliance and resistance 34

1.2.2.1 Dynamic techniques 34

1.2.2.2 Occlusion techniques 36

1.2.3 Measures of lung volume 37

1.3 The forced oscillation technique 39

1.3.1 The forced oscillation technique in older children and adults 42

1.3.1.1 Input impedance 43

1.3.1.2 Transfer impedance 45

1.3.2 The use of the forced oscillation technique in infants 46

1.3.2.1 Input impedance 46

1.3.2.2 Transfer impedance 47

1.5 The low-frequency forced oscillation technique in infants 48

1.6 Aims 48

CHAPTER 2: METHODS AND MATERIALS 51

2.1 Introduction 51

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2.2 Measurement techniques for detenriining respiratory impedance 51

2.2.1 Wave-tube device 51

2.2.1.1 Validity of the wave-tube device

51

2.2.1.2 Measurement apparatus 52

2.2.2 Pneumotachograph device 54

2.2.2.1 Spontaneously breathing patients 54

2.2.2.2 Mechanically ventilated patients 57

2.3 Forcing signals 59

2.4 Constant-phase model 61

2.5 Raised volume rapid thoracic compression technique 64

C H A P T E R 3: R E P E A T E D M E A S U R E M E N T S O F A I R W A Y A N D L U N G TISSUE

M E C H A N I C S IN R A T S 67

3.1 Summary 67

3.2 Introduction 68

3.3 Methods 69

3.3.1 Animal preparations 69

3.3.2 Measurement apparatus 70

3.3.3 Study protocol 70

3.3.4 Parameter estimation 71

3.3.5 Statistics ...71

3.4 Results 71

3.4.1 Non-invasive partitioning of Zrs 71

3.4.2 Repeated measurements in individual animals 72

3.4.3 Comparison of the lung mechanics obtained in closed and open

chest conditions 73

3.4.4 Sample numbers for non-invasive, repeated and for single invasive

studies 76

3.4.5 Mch responses detected non-invasively 77

3.5 Discussion 78

3.5.1 Validity of Pes measurement 79

3.5.2 Lung parameters 80

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3.5.3 Lung and chest wall contributions 81

3.5.4 Comparison of in situ and open chest conditions 81

3.5.5 Conclusion 82

CHAPTER 4: CHEST W A L L A N D NASAL CONTRIBUTIONS TO LOW-

FREQUENCY RESPIRATORY IMPEDANCE IN INFANTS 83

4.1 Summary 83

4.2 Introduction 84

4.3 Methods 85

4.3.1 Subjects 85

4.3.1.1 Determination of Zl and Zw 85

4.3.1.2 Characterization of Zn 86

4.3.2Measurement apparatus 86



4.3.3Study protocol and analysis 88



4.4 Results 88

4.4.1 Determination of Zl and Z w 88

4.5 Discussion 93

4.5.1 Partitioning of Zrs into Zl and Zw 94

4.5.1.1 Total respiratory impedances 94

4.5.1.2 Lung impedances 95

4.5.1.3 Contribution ofZl and Zw to Zrs 96

4.5.2 Partitioning of Zrs into Zn and Zlrs 97

4.5.2.1 Contribution ofRn to Rrs 97

4.5.2.2 Contribution of Xn to Xrs 98

4.5.3 Conclusions 99

CHAPTER 5: DEVELOPMENT OF AIRWAY A N D RESPIRATORY TISSUE

MECHANICS IN H E A L T H Y INF ANTS 101

5.1 Summary 101

5.2 Introduction 102

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5.3 Methods 1°3

5.3.1 Subjects 103


5.3.2.1 Forced oscillation technique 103

5.3.2.2 Raised volume rapid thoracic compression technique.. 104

5.3.3 Study protocol and analysis 105

5.4 RESULTS 106

5.4.1 Low-frequency respiratory system mechanics 106

5.4.2 R V R T C 107

5.4.3 Comparison between F O T and R V R T C 109

5.4.4 Factors influencing lung function parameters 109

5.5 Discussion 110

5.5.1 FOT 110

5.5.2 R V R T C :'. 112

5.5.3 Factors influencing lung function parameters 112

5.5.4 Dysanaptic growth 113

5.5.5 Comparison between FOT and R V R T C 114


CHAPTER 6: RESPIRATORY S Y S T E M I M P E D A N C E IN W H E E Z Y INFANTS. 115

6.1 Summary 115


6.3 Methods 117

6.3.1 Subjects 117


6.3.3 Study Protocol and Analysis 118

6.4 Results 120

6.5 Discussion 120


CHAPTER 7: M E T H A C H O L I N E RESPONSIVENESS IN INFANTS 125

7.1 Summary 125


7.3 Methods 127

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7.3.1 Subjects 127


7.3.2.1 Raised volume rapid thoracic compression technique.. 127

7.3.2.2 Low-frequency forced oscillation technique 128

7.3.3 Study protocol 129

7.3.4 Analysis 129

7.4 Results 129

7.4.1 Forced expiration 129

7.4.2 Low-frequency forced oscillation technique 130

7.5 Discussion 133

7.5.1 Comparison between low-frequency F O T and R V R T C 133

7.5.2 M c h responsiveness in normal compared to wheezy infants 134

7.5.3 Airway versus parenchymal responsiveness 134


C H A P T E R 8: G E N E R A L DISCUSSION 139

8.1 Methodological issues 139

8.1.1 Measurement of Zrs 139

8.1.2 Application of the constant-phase model to Zin 141

8.2 Findings of this thesis in context to the existing literature 143

8.3 Future directions 148

8.4 Conclusions

R E F E R E N C E S 151

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ACKNOWLEDGEMENTS: This thesis would not of been possible without the support and help of many people and it

really reflects the work of a dedicated group, without w h o m I could never have succeeded

in this endeavour. In particular I would like to thank the following:

The three musketeers of infant physiology; Peter Le Souei; Steve Stick and especially Peter

Sly whose knowledge I have tried to absorb and to w h o m I extend m y profound gratitude.

The respiratory fellows; Hannes Wildhaber, Mark Hayden and Nigel Dore, your mask

holding and sedation skills enabled the collection of all the data contained in this thesis.

Your friendships I will retain always.

Feri Petak and Zoltan Hantos to whom I am forever in debt. Without your expertise none

of this would have been possible.

Shane Carson; your skills as a biomedical engineer are peerless.

Celia Lanteri and Debbie Turner, whose footsteps I have had to follow, thanks for showing

m e the way.

Belaroma coffee shop; for providing the best coffee in Perth. May you long prosper.

To the endless volleyball teams I have played with, to the numerous golf partners, to the

countless drinking buddies and especially Tigger. The hours out of work made the hours

in work bearable and kept everything in perspective.

To Karen Willet for her friendship, thoughts and opinions, support and so much more. This

thesis would be a hollow shell without you.

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ABBREVIATIONS: The following abbreviations are used throughout this thesis and are usually shown in upper

case letters:

u, fluid viscosity

A cross-sectional area

C compliance

C F cystic fibrosis

C L D chronic lung disease

E elastance

FEV t forced expiratory volume (in t seconds)

FFT fast Fourier transform

F O T forced oscillation technique

F R C functional residual capacity

F V C forced vital capacity

G coefficient of tissue damping

H coefficient of tissue elastance

H G T head generator oscillation technique

I inertance

IBP infant whole-body plethysmography

ICU intensive care unit

I M integer multiple oscillation signal

Lt length

M c h methacholine

M E F V maximal expiratory flow volume

M E F maximal expiratory flows

N T M non-integer multiple oscillation signal

N S N D non-sum non-difference oscillation signal

O V W optimum ventilator waveform

P pressure

P E F V partial expiratory flow volume

R resistance

R T C rapid thoracoabdominal compression technique

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R V residual volume

S/N signal-to-noise ratio

T G V thoracic gas volume

T L C total lung capacity

V volume

V flow

V'maxFRC the maximal flow and functional residual capacity

X reactance

Z impedance

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NOTATIONS: The following notations are used in conjunction with the above abbreviations where the

abbreviation may be referred to by one or more specific regional and anatomical location

(e.g. Raw). The notations are usually shown in lower case or subscripted letters or numbers.

ao

alv

aw

aww

g

in

1

Irs

n

pl

rs

ti

(t)

tp

tr

w

airway o p e m n g

alveolar

airway

airway wall

gas compression

input

lung

lower respiratory system

nasal

pleural a

respiratory system

tissue

time (in seconds)

transpulmonary

transfer

chest wall

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STATEMENT OF CONTRIBUTION: The work contained within this thesis could not have been completed without the help of

a number of people. The author had a significant input into the design of the protocols and

methodologies and the analysis and interpretation of the data contained within this thesis.

The author was responsible for the writing of this thesis, with P.D. Sly providing

supervisory assistance on interpretation and content. The following also contributed to

sections of this thesis:

Chapter 3: This work was presented at the 1998 Thoracic Society of Australia and New

Zealand (TSANZ) and the 1997 American Thoracic Society annual scientific meetings. The

work has been published as "Repeated Measurements of Airway and Parenchymal

Mechanics in Rats B y Using Low-Frequency Oscillations' in the Journal of Applied

Physiology (1998; 85(4), 1680-1686) by Petak F., Hall G.L. and Sly P.D.. The author

contributed significantly to the design of the protocol, the analysis and interpretation of the

data and co-wrote the first and subsequent drafts of the manuscript. Other contributors are

F. Petak and P.D. Sly.

Chapter 4: The data presented in mechanically ventilated patients was part of a larger study

examining the changes in mechanics prior to- and following thoractomy. The non-invasive

pulmonary data was collected in the intact chest condition, the author was solely responsible

for the collection of this data. The author significantly contributed to the analysis and

interpretation of the data. The data describing the influence of the nasal impedance on

respiratory system impedance was presented at the 1999 Thoracic Society of Australia and

N e w Zealand ( T S A N Z ) and American Thoracic Society annual scientific meetings. The

author was responsible for the collection and analysis of the data and significantly

contributed to the interpretation of the data. F. Petak, Z. Hantos, J.H. Wildhaber and P.D.

Sly also contributed.

Chapter 5: The data in healthy infants and the relationships between oscillatory mechanical

parameters and length has been presented at the 1998 European Respiratory Society (ERS)

and 1999 T S A N Z scientific meetings. The author contributed significantly to the collection,

analysis and interpretation of the data. The multivariate analysis described in the chapter

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was carried out by K. Tiller and P. Burton of the Biostatistics department at the Institute

for Child Health Research. Z. Hantos and P.D. Sly also provided assistance in the

interpretation of the data.

Chapter 6: The author was solely responsible for the collection and analysis of the data in

asymptomatic, infants with recurrent wheeze. P.D. Sly contributed to interpretation of the

data.

Chapter 7: This data was presented in it's initial form at the 1998 TSANZ annual scientific

meeting. The author contributed significantly to the design of the protocol with assistance

from F. Petak and P.D. Sly, and was solely responsible for the analysis of the data. P.D. Sly

and Z. Hantos contributed to the interpretation of the data.

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CHAPTER 1: REVIEW OF THE CURRENT LITERATURE

1.1 Overview of respiratory physiology

The physiology of the respiratory system has been investigated, dissected and disputed.

Scientists from biochemistry, pathology, molecular biology, immunology and

pharmacology have all had input into the current understanding of the respiratory system.

The biggest advances in recent times have resulted from rapid advances in biomedical

engineering. Experiments previously never considered are n o w carried out on a daily basis.

This section outlines the basic mechanics and physiology of the respiratory airways and

tissues and the models of the respiratory system used to garner this information.

1.1.1 Basic function and mechanics of the airways

The upper airways consist of the mouth, nose, pharynx and larynx and are designed to filter

and humidify, with the larynx acting as a valve separating the upper and lower airways. The

nasopharynx represents a major source of resistance within the respiratory system [1]

particularly within young infants which are preferential nasal breathers. The main airway of

the lung, the trachea, branches into left and right main bronchi. This branching continues

down to the alveolar sacs, with the total cross-sectional area increasing with each

generation. The initial 16 generations (trachea to terminal bronchioles) are known as the

conducting zone, the subsequent 3 (respiratory bronchioles) generations are the transitional

zone, with the final 3 (alveolar ducts and sacs) termed the respiratory zone. The large

trachea and main bronchi are supported by U-shaped cartilage which are joined posteriorly

by smooth muscle bands. Further down the airway tree (generation 2-11) this cartilaginous

support changes to irregularly shaped, helical plates. Beyond this point the airways are

directly embedded within the pulmonary parenchyma and there is no cartilage within the

airway walls. Instead, the caliber of the airway is maintained by the elastic recoil of the lung

and hence lung volume.

Within the conducting airways, gas flows by bulk flow. The resistance of these airways is

determined by the geometry of the airway and the flow of the gas passing through the

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airways. In the transition and respiratory zones the bulk movement of the gas ceases, with

movement of the respiratory gases (0 2 and C 0 2 ) occurring by diffusion. The resistance of

these airways is also determined by the geometry and the flow of the gas. The geometry of

the intra-parenchymal airways is related to the localised elastic recoil pressure which is in

turn proportional to lung volume. In addition, flow is determined by diffusion and hence is

relatively small. Thus within the conducting and respiratory zones the predominant influence

on the resistance of the airways is lung volume. Most of the flow resistance in the intra

thoracic airways occurs in these lower regions due to the proportionally high total cross-

sectional area of these generations [1].

The inertance of the airways can be considered to be negligible during quiet breathing, but

at higher frequencies (>10 Hz) plays an important role. As inertance is directly proportional

to the acceleration of the gas (see Section 1.1.3.3 below) and as the intra-parenchymal

airways conduct flow by difilision, inertance will be primarily determined by the central and

upper airways.

The resistive and inertive mechanical properties of the conducting airways are independent

of frequency, as shown in a number of studies, in dogs [2,3], cats [4], rats [5] and in

humans [6,7].

1.1.2 Basic function and mechanics of the lung parenchyma

The lung can be grossly divided into two major functional components: lung parenchyma

containing the gas exchange tissue of the respiratory zone and the non-parenchyma

comprising the conduction structures, airways, blood vessels and the coarser connective

tissue components. The mechanical properties of the lung parenchyma are in part defined

by the connective tissues that are embedded within the lung tissues. This network of

connective tissue arises from a number of sources: 1) peripheral fibres originating from the

pleura and penetrating into the inter-segmental and interlobular septa; 2) axial fibres,

consisting of tissue sheaths extending out from the airway tree into the acini region,

becoming the fibrous network surrounding the alveolar ducts and; 3) the alveolar septal

fibers which join axial and peripheral fibre networks and interlaces with the capillaries [8,9].

The other major factor affecting the parenchymal mechanics is the large air-liquid interface

21

lined with surfactant controlling surface tension [9]. These two factors can be loosely

described as the tissue and surface mechanical factors. The manner in which they interact

defines the mechanical properties of the lung parenchyma.

The elements of the peripheral fibre network increase with the cube root of lung volume,

while at a fixed lung volume this system is relatively independent of increasing surface

tension [8]. The axial fibre network consists of force-bearing elements forming the helical

structure of the alveolar ducts, while the septal tissue network of the alveolar walls links

these force-bearing elements and the rims of the adjacent alveolar ducts. This network is

extended by the surface tension acting on the alveolar walls, as surface tension increases,

the ducts widen and the alveolar walls shorten by rearrangement of the septal tissues [8].

The alveolar surface lining layer (surfactant) regulates the mechanical forces that arise at

the air-liquid interface. The surfactant layer reduces the acting surface tension of the alveoli,

hence increasing the elasticity of the lung and reducing the work of breathing. In addition

the surfactant layer also tends to increase the stability of the alveoli and keep the alveoli dry.

The interaction between surface and tissue forces depends on the volume history of the

lung, for instance at functional residual capacity (FRC) following an inflation to total lung

capacity (TLC) the surface tension approaches zero and hence the elastic recoil of the lung

is dominated by the tissue forces [10]. However, following prolonged tidal breathing,

surface tension increases and volume decreases causing surface forces to contribute

increasingly to the elastic recoil.

The mechanical properties of the respiratory tissues are highly dependent upon the

frequency at which the measurements are made. This frequency dependence of the

respiratory tissue mechanics is well established with work in both animal [11-15] and

human studies [7,16-21].

1.1.3 Modelling the respiratory system

In order to study respiratory mechanics the system under investigation must first be defined,

allowing the behaviour of the system to be investigated under a number of conditions. The

advancement of mathematics, physics and engineering has allowed the behaviour of gas,

liquid and solid substances under a variety of physical conditions to be defined. These

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definitions have been used by physiologists to further characterize the behaviour of the

respiratory system. The use of mechanical and electrical modelling has allowed the naturally

dynamic state of the respiratory system to be investigated in further detail. Any system can

be described as a collection of elements, whether the system be the upper airway, lung

parenchyma, chest wall or the complete respiratory system. These elements are described

according to their use or storage of energy and are called resistances, elastances and

inertances. Respectively these three elements handle energy in a unique way: dissipation of

energy by way of friction (viscous), storage of energy using potential energy (gravitation

or elastic) and storage of energy by means of kinetic energy [22].

In respiratory mechanics these elements are characterized in terms of pressure changes

across an element (AP) and the corresponding flow ( V ) passing through the element. The

relationship depends on the element's geometry and physical characteristics and is called

the constitutive relation and is defined as:

P = f ( V ) -(1)

1.1.3.1 Resistance

A n ideal fluid resistor exhibits a decrease in pressure directly proportional to flow such that:

P = RV(t) -(2)

where: R=resistance and t=time

The simplest version assumes a low flow-pressure relationship, with a steady laminar flow

at a low Reynolds number (exhibits a parabolic velocity profile) in a straight tube of length

(L) and cross-sectional area (A) containing a fluid of known viscosity (|i) and hence

resistance of the tube, of radius (r) is given by Poiseuille's formulae [22]:

R = 8 |iL /TC r4 -(3)

In the presence of bias flows or at higher frequencies the resistance may become dependent

on other factors such as frequency (f) or density (p). In this case resistance can be

approximated as:

R « R D C + L(V2 u / TC r4)(r2 co /v)1/2 -(4)

where: R D C is the resistance calculated from the steady state flow-pressure relationship, v

is the kinematic viscosity (u/p), © is the angular velocity (27if).

23

If the flow is passing through an orifice then the pressure drop across the orifice is

proportional to the square of flow and hence resistance may be given as [22]:

R = |V|p/(2Cd2Ao 2) -(5)

where: Cd is the discharge coefficient and A o is the orifice area.

Thus resistance is a complex relationship between flow and pressure, the nature of which

will be defined by the length and radius of the tube and the viscosity and density of the fluid

passing through it. In the case of a distensible cavity with viscous walls (such as found in

the parenchyma), resistance may be described as a function of the geometry of the unit

involved and its visco-elastic moduli. Resistive elements dissipate heat proportional to the

friction or viscance of the fluid within the tube [22].

1.1.3.2 Elastance

A n ideal fluid elastance (E) exhibits pressure changes in proportion to volume (V). The rate

at which volume changes is the volumetric flow rate and so:

P(t) = EV(t) -(6)

Within the respiratory system the most important source of fluid elasticity is gas

compressibility. This represents a major source of elastance at higher frequencies which

tend towards an adiabatic condition [22]. In this case elastance is a linear function expressed

as:

E = pPo/V0 -(7)

where: P is unity if gas compression is isothermic and is the ratio of specific heats if it is

adiabatic, V 0 is an ideal gas volume at an absolute mean pressure Po.

The behaviour of a solid elastance is dependent on the geometry of the object and its elastic

moduli. Within the respiratory system solid elastic elements tend to be dependent on volume

history and are profoundly non-linear. Solid elastances may be linearized to a reference

volume. Hence the tangent slope at a particular point on the pressure-volume curve is the

elastance at that volume and is termed the incremental modulus [22]. Within the respiratory

system airway walls, chest wall and lung parenchyma are examples of solid elastances.

24

1.1.3.3 Inertance

In the absence of viscosity, the pressure or force (F) applied across a tube of area A and

length L is proportional to the acceleration (du/dt) and the mass (pLA) of the fluid in

accordance with Newton's law.

F = /2PLA (du/dt) -(8)

where: n = 1 if the velocity profile is blunt and 4/3 if parabolic, p is the fluid density, u is

the fluid velocity and d/dt represents a rate of change with respect to time t.

At higher frequencies and in large ducts (or airways) the velocity profile will tend towards

a blunted aspect. The ideal elemental equation for inertance (I) relating flow to pressure is:

P(t) = I(d/dt)V(t) -(9)

where: I = «pL/A -(10)

It can be seen from the ideal elemental equation that inertance is largest during rapid flow

changes and zero during periods of constant flow. For a solid object of mass M and area

A, the solid inertance can be given as: I = M / A 2 -(11)

In this instance flow represents the rate of volume passing around the surface of the solid

object. Within the respiratory system the motion of the chest wall or airway dilations are

examples of solid inertances [22].

1.1.3.4 Models of the respiratory system

The elements described above can be combined to produce a model that will attempt to

explain the behaviour of the respiratory mechanics under certain defined conditions. The

remainder of this section outlines some of the more common models of the lung, their

advantages and limitations.

A mechanical model of the lung was initially introduced by Otis etal [23]. It allowed for

a single compartment of constant elastance attached to a constant flow resistor. It assumed

that lung volume and flow did not affect the mechanical properties of the respiratory system

and that inertial properties were negligible and was represented as:

P(t) = EV(t) + RV(t) -(12)

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While these assumptions are an oversimplification of the respiratory system, it is a

reasonable assumption for quiet breathing with small tidal volumes. Initial techniques

utilizing this or similar models were restricted by the available technology but recent

advances in computer technology have allowed the multiple regression analysis technique

to be applied to the model described above. Under dynamic conditions, the inertial

properties of the respiratory system should be allowed for and hence the lung can be

modelled as:

P(t) = EV(t) + RV(t) + IV"(t) -(13)

All models of the lung assume linearity. That is, the resistive, elastic and inertive

components do not alter under different measurement conditions. In nature the respiratory

system is dependent on volume history, volume and flow. Similarly the volume-pressure

relationship of the respiratory system in not linear, but rather sigmoid. If the limitations of

the models used are understood and the conditions under which measurements are made

kept within certain, known boundaries, then linear models of the lung can provide valuable

information.

If overdistention occurs, for example in ventilated patients with lung disease, then the model

used to track respiratory mechanics may require a volume-dependent elastic component

[24]. This volume dependence of the respiratory system can be accounted for by including

an extra component in the above equation such that:

P(t) = (El + E2V)V(t) + RV'(t) + IV"(t) -(14)

where: El and E 2 are the non-volume and volume dependent components, respectively.

The use of this model may improve the interpretation of pressure-volume data even if there

is no evidence of over-distention [25].

A time-varying sinusoidal signal (such as tidal breathing) can be written as a function of

frequency rather than time. If w e examine an idealized tidal breathing trace (represented as

a sinusoid in Figure 1.1 below), the time taken to complete a single breath is called the

period (T) of the signal. The number of breaths per second is termed the frequency (f) and

26

is the reciprocal of the period (T = 1/f). The breathing frequency can also be given as:

(D=27rf - 0 5 )

where co is the angular frequency.

Any periodic signal can be written as a vector with a corresponding phase angle (9).

Giving an equation: X = Bcos(cot), Y = Bsin(cot) -(16)

and cos(oot) = sin(cot + 7t/2)

where: X and Y represent the real and imaginary components, respectively; B is the

magnitude of the signal at a given point and (cot) represents the angle 0 at that point.

Figure 1.1: Idealized tidal breathing trace. A m p , amplitude; T, period; oo, angular frequency.

If the real component (X) represents that portion in phase in flow, then the imaginary

component (Y) represents that portion out of phase with flow, where the phase difference

is 90°. If w e then equate this to a point in the complex plane, w e could replace these two

equations with a single equation of motion such that:

W = X + jY = A m p (cos(oot) + j sin(ot)) -(17)

= A m p eJ<at

where: W is a complex number and j = V-l

Hence if equation 17 is rewritten in terms of the variation in flow at the airway opening can

be written as:

V(t) = V(oo) cos(cot) = V'(co) e""* -(18)

Similarly, pressure can be written as:

Pao(t) =Pao(co) cos(cot+0rs) = Pao(oo)ej(<Bt + 0is) -(19)

where Grs accounts for the phase shift of the pressure wave relative to the flow.

27

The impedance (Z) can hence be written as the complex relationship between flow and

pressure and is represented as:

Zrs = Pao(oo) <P*+** / V(a>) e*01

= (Pao(oo)/V(a)))eie,B

= |Zrs|eiere -(20)

where |Zrs| is the magnitude of the pressure difference relative to the flow and a phase angle

9rs indicating the phase angle of the pressure difference relative to flow [22].

Thus a signal of any kind can be represented as an amplitude and a phase angle or

frequency. This can be extended to a signal that is a combination of a number of individual

frequencies as represented below

f(x) = Bi * > " + B 2 e^24 + B 3 J*

31 + .... -(21)

This form of representation is termed a Fourier series. By breaking a complex signal into

its composite frequencies it can be analyzed more efficiently, producing a graph that plots

magnitude on the y axis and frequency on the x axis. So by using a Fourier transform, a time

varying signal applied at the airway opening can be decomposed into its constitutive

frequencies and the flows and pressures at those frequencies analyzed to produce a

respiratory system impedance.

The mechanics of the respiratory tissues are frequency dependent. The ratio of pressure to

flow at a frequency oo is the respiratory impedance (Z), such that:

Z = P ( ( D ) / V ( G ) ) -(22)

Impedance is a generalization of resistance, but whereas resistance describes only resistive

(or frictional) induced pressure differences, impedance describes pressure differences across

resistive, elastic or inertive elements [22]. The respiratory system impedance (Zrs) can be

divided into its real (resistive, Rrs) and imaginary (reactive, Xrs) components:

Zrs = Rrs+jXrs -(23)

The real component is that portion of the impedance in which the pressure changes are in

28

phase with the changes in flow, whereas the imaginary component is the part of the

impedance which is out of phase with flow. If this approach is applied to the R-I-C models

described above then impedance can be given as:

Z = R + j © I - l / © C -(24)

This model can be applied to each of the components of the respiratory system, providing

their pressures and flows can be determined [22]. As described above the mechanics of the

respiratory system are frequency dependent, however this frequency dependence differs for

Z = R Z=j©I Z = -j(E/G>)

Figure 1.2 Resistance-Inertance-Compliance model.

the various components of the respiratory system (airways versus parenchyma) (See Section

1.1.1 & 1.1.2). At low frequencies ( < 2-4 Hz) the respiratory tissues dominate the

resistance and reactance of Zrs, but with increasing frequency this contribution decreases.

At frequencies above 10 H z the mechanical properties of the airways account for the

majority of Zrs. It is possible to separately determine the airway and tissue mechanical

properties of the respiratory system by the use of a model that includes both airway and

tissue compartments. One such model is the constant-phase model of the lung which has

an airway compartment containing a resistance and inertance, and a viscoelastic tissue

compartment that exhibits a stress relaxation response with a pressure decay described as

a power-law in time. [3,4,26]. This constant-phase model (CPM) of the lung has two tissue

properties: tissue damping (G) and tissue elastance (H) and these relate to Rt and Et as

shown:

Rt = G/©a, Et = H(©/©a) -(25)

where: a = (2/7t)tan"1(H/G)

Hence the constant-phase model of the lung can be given as:

29

Z = R + j©I + (G - jH)/©a -(26)

The CPM of the lung allows the simultaneous assessment of the airway and tissue

components without the use of more invasive techniques, such as alveolar capsules. T o do

this the model must make certain assumptions. The lung is assumed to be a linear system

with a single airway resistance and inertance and a homogeneous tissue compartment. The

assumptions of the model will only remain valid under these conditions. Should peripheral

airway inhomogeneities or regional tissue changes occurs then the model may no longer

produce a valid separation of the airway and tissue compartments. Lutchen etal.[27] used

a model to investigate the role of airway inhomogeneities, airway wall shunting and tissue

viscoelasticity on the frequency dependence of the respiratory tissues, in addition the

authors examine the role of lung inhomogeneities and airway wall shunting to the estimated

airway and tissue parameters derived from Zrs. The authors simulated various levels of

airway constriction and showed that under conditions of severe constriction the frequency

dependence of the tissues increased and that the C P M produced an overestimation of the

tissue damping and hysteretic properties of the respiratory tissues [27]. These results have

been confirmed in animal studies [5,28]. Using alveolar capsules Lutchen et al [28] showed

that separation of airway and tissue properties using the C P M were reasonably consistent

with those values derived from alveolar capsules, even under conditions of moderate

constriction. A further study by Petak and coworkers [5] used an inert gas to demonstrate

that the C P M was able to accurately separate the airway and tissue compartments of the

lung under normal and moderately constricted conditions. Under severe constriction the

authors showed an increase in tissue damping not mirrored by changes in H, a decrease in

inertance and large increases in hysteresivity. These changes were attributed to peripheral

airway inhomogeneities [5]. Thus the constant-phase model allows the non-invasive

separation of Zrs into airway and tissue mechanical properties. Under normal conditions and

during moderate constriction the model has been shown to reliably partition Zrs. However,

if airway inhomogeneities occur, such as following severe constriction then the model may

not longer accurately represent the mechanical properties of the respiratory system.

The first model put forward taking into account the frequency dependence of the

respiratory system was that of DuBois [29] (Figure 1.3). This model can be applied to either

30

input or transfer impedance data (see Section 1.3.1) and as data from Ztr techniques are

influenced by gas compressibility it is the most appropriate model for this method [29].

The models described above assume a single airway resistance and are termed lumped

parameter models. Inherent in their assumptions are that the airways are rigid, and that

ventilatory patterns are homogeneous to all regions of the lung. Lumped parameter models

are limited if the distribution of flows and pressures in the upper and lower airways is of

interest. If the physical properties of the airway tree (such as airway geometry, wall

properties or branching) are to be investigated or if the oscillatory frequencies are high

enough that standing wave phenomena occur in the airways. In these cases a distributed

model of the lung should be used [22].

Z a w Zti

Vao Vbs

— Pb = 0

Figure 1.3: Electrical equivalent of the DuBois model. Zaw, Zt and Zg represent the impedances of the airways,

tissues and alveolar gas compartments, respectively. The forcing signal may be applied at the airway opening

(Pao) or at the body surface (Pbs).

31

A distributed model of the lung takes into account the spatial dimensions and branching

topology of the airway tree. Pressures and flows are calculated not at access points such

as the airway opening or at the alveoli but at every airway position between the two. In

addition, the parenchyma is not considered one compartment, but rather a widespread organ

with possible differing pressures and volumes throughout [22]. The mechanics of a section

of airway, of length dx, can be described by a series impedance of resistance and inertance

in parallel with a gas compliance, such as due to compression (Cg), gas thermal

conductance and airway wall distensions (Iaww, R a w w and E a w w ) and includes the

mechanical support of the surrounding lung parenchyma.

The distributed model of airway mechanics requires certain assumptions to be valid: the

pressures at the junctions between airways must be equal and net flow into an airway

bifurcation must be zero. Both conditions can be met if low-amplitude flow oscillatory

signals are used [22]. Once the impedance in a segment is calculated, recursive analysis is

used to calculate the mechanics of the whole airway tree and by using a distributed model

the mechanics at a given airway generation can be determined. These models however,

assume constant mechanical properties of all airways within a generation. A stochastic

model allows for a range of airway diameters and lengths within an airway generation [30].

A n extension of this model which allows for varying airway wall thickness has been used

to demonstrate that the variability in airway responsiveness can be replicated when the

distribution of airway properties within the lung are taken into account [30]. A distributed

model has also been applied to the non-linear stress-strain characteristics of the lung

parenchyma by Maksym and Bates [31]. The authors introduced a model accounting for the

mechanical properties of elastin and collagen fibres within a canine lung tissue strip and for

pressure-volume curves obtained from normal, emphysematous and fibrotic lungs.

1.2 Measurement of respiratory mechanics in infants

During the first two years of life the respiratory system undergoes profound developmental

changes, which until recently were not clearly understood. Over the past ten years the

number of techniques available to monitor respiratory mechanics in infants has greatly

improved, increasing the knowledge of both normal developmental changes and progression

of disease in this age group.

32

The joint committee of the A T S assembly on Paediatrics and the E R S Paediatric assembly

categorized these infant pulmonary function tests into the following areas: measures of

forced expiratory flow, measures of compliance and resistance, measures of lung volume

and other tests (including measures of tidal volume, respiratory patterns, chest wall motion

and gas mixing)[32]. This thesis concentrates upon the developmental aspects of the low-

frequency forced oscillation technique (FOT) as it pertains to infant lung function. Those

techniques currently available will be briefly reviewed, followed by a detailed account of the

F O T , its use in older children and adults, the basic underlying assumptions, interpretations

and applications surrounding its use.

1.2.1 Measures of forced expiratory flow

The maximal expiratory flow-volume relationship ( M E F V ) has played a central role in the

evaluation of the function of the intra-thoracic airways [32]. Within older children and

adults, the measurement of M E F V via spirometry is the standard lung function test.

The investigation of flow-volume relationships in infants has lead to the development of a

number of techniques. While these tools have contributed much in the way of valuable

information about the developing respiratory system there exists a number of controversies

surrounding the techniques and their interpretation [33,34].

The forced deflation technique was initially carried out in 1977 by Motoyama [35] and

involves inflating the lungs to a raised volume followed by the application of a negative

pressure. The negative pressure is applied until residual volume (RV) is reached or to a

maximum of 3 seconds and results in increased flows over the whole volume range. The

procedure is repeated until 3 identical M E F V curves are acquired and from these,

measurements of forced vital capacity (FVC) and forced expiratory volumes (FEVt)can be

obtained. Maximal expiratory flows (MEF) at volumes equating to 1 0 % (MEFio) and 2 5 %

(MEF2 5) of F V C from R V are determined and the ratios MEF10/FVC and M E F 2 5 / F V C

calculated which are thought to represent an index of the upstream airway segment [36].

While published data from the technique are limited, the results appear to be highly

reproducible. This technique is limited to intubated patients and thus its widespread

33

application as a diagnostic tool in infant lung function is doubtful.

The rapid thoracic compression technique (RTC) [37] is a common technique for assessing

lung function in infants. Baseline studies have been conducted in healthy infants [38], infants

with acute and chronic lung disease [39] and infants with cystic fibrosis (CF) [40]. In

addition, airway reactivity using dilator and constrictor agents has been examined [41-43].

For the R T C technique infants are sedated and a sudden pressure is applied to the thorax

at end-tidal inspiration using a positive pressure source attached to an inflatable jacket [44].

The resultant forced expiratory flows are plotted to produce partial expiratory flow-volume

(PEFV) curves. Prior to the R T C manoeuvre, F R C is determined as an identifiable end-

expiratory point and maximal flow at F R C (VmaxFRC) is then reported. The R T C

technique is surrounded in controversy centred around the active control of end-expiratory

lung volume in infants and the question of whether or not flow limitation has been achieved

during the R T C manoeuvre [32]. F R C is known to vary with dead space, sleep state and

disease state [45] and this variation may be responsible for the high variability of V m a x F R C

in normal infants [44]. Further variability in V m a x F R C may result from the possible lack

of flow limitation during the R T C manoeuvre [32]. While flow limitation has been achieved

in infants with bronchopulmonary dysplasia (BPD) or severe airway obstruction [32],

information regarding flow limitation in normal infants is lacking.

The flow-volume relationship obtained with the RTC technique is contained solely within

the tidal volume range. In adults and older children, the use of an extended volume range

has been shown to produce more useful measures of forced expiration than those produced

from the tidal volume range. This principle has been applied to infants using the raised

volume adaptation of the R T C method ( R V R T C ) [46,47], thus allowing measurements of

FEV t to be determined. The infant's lungs are raised to a standard inflation pressure (20

c m H 2 0 ) , with the jacket compression force also standardized to transmit a further increase

of 20 cirdHfeO to the airway opening [48]. This standardization of transmission pressure

results in pressure-independent forced expiratory flows and ensures that comparisons

between individual infants and laboratories can be made [48]. While this technique avoids

many of the limitations associated with the R T C method it does not ensure that flow

limitation occurs. It could be argued that as the volume to which the infants lungs are raised

34

and the applied compressive pressure is standardized, then it is not necessary to produce

flow limitation. Further work into this question is still needed. While little data exists on the

ability of the R V R T C method to determine the difference between normal and diseased

groups it appears that the technique will provide better assessments of airway function that

the R T C method.

The RTC and RVRTC methods allow non-invasive measurement of lung function in infants

and as such have a wider application than the forced deflation technique. Overall, these tests

can only provide information on limited changes within the lung. Information is not

available on the behaviour of the airways or the tissues separately and as such these

techniques can only provide a finite amount of information on the development of

parenchymal disease or the relative growth patterns of the airways and pulmonary tissues

in normal development.

1.2.2 Measures of compliance and resistance

The number of techniques developed to measure compliance and resistance has grown

rapidly. These techniques can be split into two approaches: firstly, measurement of lung

mechanics during a number of complete cycles of spontaneous breathing or mechanical

ventilation and secondly, the determination of respiratory system resistance (Rrs) and

compliance (Crs) during controlled flow conditions, such as relaxed expiration or zero flow.

These are termed dynamic and occlusion techniques, respectively. While the application of

techniques and interpretation of results differ, a number of important clinical and

physiological investigations have been carried out. The underlying assumptions and

limitations of the main techniques are reviewed, as well as the information that can be

obtained:

1.2.2.1 Dynamic techniques

The dynamic mechanics of the respiratory system describe the mechanical behaviour of the

chest wall, lungs and airways throughout the breathing cycle or, less commonly, during

mechanical ventilation. The assessment of dynamic mechanics relies on the measurement

of pressure cycles and changes in flow and volume that these pressures produce [49].

Dynamic mechanics have been assessed through a variety of models or in combination with

35

other methods, such as plethysmography, forced oscillation and the interrupter technique.

The background, underlying assumptions and limitations of the various analytical models

used in this technique have been outlined previously (Section 1.1.3.4) and the forced

oscillation technique is described in detail in Section 1.3.

Whole-body plethysmography was first described in adults in 1956 [50] and was rapidly

customized for use in an infant population [51]. Infant whole-body plethysmography (IBP)

is an important tool for the simultaneous assessment of lung volume and airway resistance.

For this technique the infant is placed inside a closed chamber and breathes through a

pneumotachograph, allowing tidal flow and hence tidal volume to be determined. Following

a brief occlusion of the airway opening, changes in airway pressure (Pao) and hence

alveolar pressure (Palv) can be assessed. These changes allow thoracic gas volume ( T G V )

to be calculated. Airway resistance is determined by keeping the respired gas saturated at

body temperature and pressure (BTPS conditions), and relating changes in Palv to

concurrent changes in flow. Regression equations for R a w have been reported in infants up

to two years of age [32,52]. The values have been found to vary with race, gestational age

and the flow at which the measurements are determined. Disadvantages of the IBF

technique is that it requires complex equipment and considerable training is needed in order

to produce reliable results. In addition a number of assumptions must be made, most of

which apply to the measurement of T G V and have hence been detailed in Section 1.2.3.

Overall, the use of the IBP technique has been limited and while useful information has been

produced, it is not suitable for bedside investigations and is unlikely to obtain widespread

use as a clinically acceptable technique for infant lung function [32].

The interrupter technique is a useful method for partitioning the respiratory resistance into

airway and tissue components. W h e n the airway is occluded during expiration there is an

abrupt increase in pressure (Pinit) followed by a slower, secondary rise in pressure to a

plateau. The difference between the total change in pressure and Pinit is Pdiff and represents

the tissue visco-elastic component [53]. If the correct equipment and analysis is used, then

reproducible and reliable estimates of airway and lung tissue mechanics can be

determined. [54]. However, investigators have shown that errors can be introduced by

unsupported upper airways [54], volume history [55,56] and by ventilation inequality that

36

may be present in infants with parenchymal disease [57]. While studies have been conducted

in dogs [56], cats [55,58] and rodents [59], in humans most studies have been conducted

within the intensive care environment [60,61]. Recently the technique has been applied to

awake healthy children [62-64]. However the technique has been limited to that of a

research tool and is unsuitable for use in healthy infants.

1.2.2.2 Occlusion techniques

While dynamic techniques have supplied a wealth of information regarding the mechanical

behaviour of the airways, lung and chest wall in patients with respiratory disease, those

studies examining the properties of healthy infants have only been able to recruit small

numbers. Less invasive occlusion techniques have shown themselves to be useful in

obtaining data in healthy infants. The measurements however, reflect the entire respiratory

system. In young infants, the chest wall is highly compliant relative to the normal lung [65].

This changes as the infant grows, with the relative contribution of the chest wall increasing.

B y two years of age, the contributions of the chest wall and lung to the respiratory system

compliance are approximately equal [65]. The underlying assumption for all occlusion

techniques is the complete relaxation of the respiratory muscles. In most cases this is done

by invoking the Hering-Breuer reflex [66]: the infant's lungs are held above F R C causing

a stimulation of stretch receptors and an inhibition of inspiratory effort, thus prolonging the

expiratory pause and relaxing the respiratory muscles [67]. During these periods of

relaxation, Pao will reflect the elastic recoil pressure of the respiratory system allowing Rrs

and Crs to be calculated [32]. The most c o m m o n occlusion techniques are the occlusion

technique itself and passive flow-volume and weighted spirometers. The occlusion

technique also has a number of adaptations. These are the multiple occlusion, single-breath,

multiple interruption and expiratory volume clamping techniques.

Occlusion techniques rely on the assumption that during airway occlusion Pao reflects

elastic recoil. This may not be the case if the respiratory muscles are not completely relaxed,

or pressure throughout the respiratory system has not equilibrated [32,67]. Equilibration

may be a problem in infants with severe airway obstruction [68]. Results may also be

influenced by a change in the state of the patient or the measurement conditions. These may

include sleep state, health of the infant, changing position or addition of pharmacological

37

agents that cause changes in the respiratory system (such as constrictors, bronchodilators

or sedatives). Currently, the techniques used have not been standardized and as such

comparison of results is difficult. Passive respiratory mechanics are non-invasive, simple to

use and well tolerated by most infants. The techniques have been used to assess the

influences of smoking, growth and gender in a population of healthy infants [69]. Crs has

been examined in healthy preterm infants [70] and in ventilated infants [71,72] as well as

a variety of other investigations [73-75]. The technique can be applied to both

spontaneously breathing and mechanically ventilated subjects. The major limitation is that

the techniques cannot separately determine the properties of the airways and lung tissues.

In neonates and young infants where the majority of disease is parenchymal in origin this

may play a major role in a decision on which technique should be used.

1.2.3 Measures of lung volume

The volume dependence of the respiratory mechanics is well established [76-79]. Hence it

is important to have a volume landmark with which to compare the mechanical

characteristics of the respiratory system. In infants the most commonly measured lung

volume is F R C [32]. This section outlines the three most widely used techniques to

determine FRC: whole-body plethysmography, helium dilution and nitrogen washout.

Infant whole-body plethysmography has allowed investigators to simultaneously determine

R a w and thoracic gas volume ( T G V ) (as discussed in Section 1.2.2.1). The technique has

the advantage that it determines all gas within the respiratory system, thus allowing an

assessment of the amount of gas trapping in obstructive lung disease [80]. Four major

assumptions must be made in using IBP to determine T G V [81]. These are: 1) there is no

flow in the airway during occlusion, 2) pressure changes applied to the lung are

homogeneous within the pleural space, 3) pressure-volume changes are limited to the

volume of gas within the thorax, and 4) changes in pressure and volume are isothermal.

If the upper airway acts as a shunt during airway occlusion, then flow may occur hence

leading to an overestimation of T G V [82,83]. Eber et al [84] have shown that unequal

alveolar pressures occurring due to significant airway obstruction may cause an

overestimation of T G V in wheezy infants. If the occlusion is activated at end-inspiration,

38

and thus higher lung volumes, then the pressure equalization time between areas of the lung

will decrease, hence minimizing the potential overestimation [85]. During disease, regions

of the lung with high resistance and elastance may act as rigid spheres and gas trapped

within these spaces may not be subjected to rarefaction or compression. Under- or

overestimation of T G V in infants recovering from bronchiolitis [86] or airway obstruction

[87] has been reported. The volume of gas located within the gastrointestinal tract would

be included in an estimate of T G V , however, this volume is thought to be insignificant.

While this has been confirmed in adults [88] and in healthy and sick infants [68] it may be

possible that under conditions of increased abdominal gas such as in patients with cystic

fibrosis (CF) there is potential for the introduction of errors to the measurement of T G V

[81]. The construction of the plethysmographic chamber will influence transmission of heat

across the walls of the chamber. If the chamber is calibrated to the approximate rate of

respiratory efforts and size of the chamber chosen carefully then introduced errors can be

compensated for [32]. However, the highly variable respiratory rate in infants that may

occur during a test, or following the administration of inhaled agonists will continue to

provide a source of error. While reference values have been derived for F R C in infants, IBP

is limited due to the expense of equipment and the considerable operator training required.

The most common method for measuring FRC in infants is the helium dilution technique

[89]. The technique has been applied to infants with bronchiolitis, C F and respiratory

distress syndrome and uses the gas equilibration principle occurring between an unknown

gas volume and a known gas volume in communication with each other [32]. The

equipment used is relatively inexpensive and can be applied in outpatient and I C U settings.

The determination of F R C assumes that equilibrium occurs between all lung units and the

airway opening. If gas trapping is present due to severe obstructive airway disease then the

measured F R C may be underestimated. Underestimation may also occur if the inspired

oxygen is high [32].

The nitrogen washout technique measures the volume of nitrogen washed out of the lung

during a period of time as the infant re-breathes from a nitrogen free gas source. The

required equipment is less complex than that required for plethysmographic or helium

dilution techniques. If a mass spectrometer is used, then an open circuit system without a

39

collection bag can be used [32]. Errors can be introduced due to analyzer response time,

lag time between flow and gas concentration and sampling rate [32]. In addition,

corrections must be made for nitrogen originating from the tissues and blood, which

typically can be up to 5 % [90], but may be larger in infants with lung disease. The technique

has been used to provide information on infants with chronic lung disease (CLD) [91],

respiratory failure [92] and in mechanically ventilated infants [93].

1.3 The forced oscillation technique

The use of the forced oscillation technique (FOT) to predict the mechanical properties of

the respiratory system was first carried out by DuBois et al. in 1956 [29]. The authors

concluded the airways could be characterized as a resistance-capacitance system, leading

to the chest wall and diaphragm which was suggested to be represented as visco-elastic

massive surfaces. This technique has gone from the gross clinical tool described by DuBois

to one able to provide specific insight into the physiology and mechanical behaviour of the

respiratory system [94]. In its most basic form, the F O T uses small time-varying forces to

perturb the respiratory system. The responses deduced from these perturbations give

information about the structural and mechanical properties of the respiratory system [22].

Due to the flow and volume dependence of the respiratory airways and tissues (see Section

1.1.1 and 1.1.2 above), the construction of the oscillatory signal is of crucial importance.

The simplest way of determining the frequency dependence of the respiratory system

mechanics is the successive application of sinusoids of different frequencies. However, this

is a time-consuming procedure susceptible to changes in patient condition. The use of

known signals and appropriate signal processing allows simultaneous determination of the

Zrs across a defined frequency range. These signals can take three forms: 1) random noise;

2) a train of impulses and; 3) composite signals. Random noise is disadvantaged due to the

'noise' of the respiratory system (e.g. spontaneous breathing or inadequate respiratory

muscle relaxation) hence resulting in a low signal to noise (S/N) ratio. The use of composite

signals allows the total signal energy to be maximized for a given amplitude limit and for

the S/N ratio to be optimal across the whole frequency range [95]. This can be carried out

by weighting the amplitude of the signal towards the lower frequencies and assigning the

phase of each frequency component such that the peak-to-peak amplitude of the signal is

40

minimized [95]. This optimized signal can then be applied to spontaneously breathing

patients without discomfort [95], or if an apnoeic period is induced, the low-frequency Zrs

spectra determined [96]. Another approach is to create the signal such that it delivers

sufficient volume to maintain gas exchange, while minimizing the peak-to-peak Pao, termed

optimal ventilation waveforms ( O V W ) [97] (See Section 1.3.1 below).

The second consideration in signal construction is the frequency components to use. As the

parenchymal mechanics exhibit nonlinear behaviour (e.g. volume dependent behaviour) the

respiratory impedance will be a function of both the linear (frequency dependence) and

nonlinear (volume dependence) behaviour of the lung tissue mechanics [98]. It is thus

desirable to minimize the influence of these non-linearities. One approach is to construct a

signal where the S/N ratio is maximized and the peak-to-peak pressure amplitude and flow

minimized [99]. Properly constructed, small amplitude oscillations should not affect the

respiratory mechanics. Peslin et al. [100] demonstrated that in ventilated rabbits the

oscillatory signal did not affect either baseline lung mechanics or the lung response to

histamine. If the measured system is nonlinear, then output harmonics can be produced that

match the periodicity of the input signal and thus the measured impedance will be a function

of the spectral characteristics of the input signal used [98]. The presence of non-linearities

may not be detected by the coherence function (y2) when using composite signals, as y2 will

only differ from unity in the presence of extraneous noise [99]. A second approach to

reduce the effects of distortion is the use of periodic signals with a selection of frequency

components that are not multiples of the fundamental harmonic [99]. This introduction of

non-integer multiple signals (NTM) reduces the effects of harmonic distortion within the

system as the nonlinear harmonics produced by the system will not interfere with frequency

components in the forcing signal. Using a nonlinear model of the lung, Daroczy et al [99]

demonstrated that the N T M signal leads to a significant improvement in impedance

estimation when compared to a standard composite signal. However, M M signals are still

subject to nonlinear distortions associated with inter-modulation between frequency

components of the input (termed 'cross-talk') [98]. Cross-talk occurs when some of the

energy component at a frequency is folded back onto a lower frequency, so as to cause the

energy magnitude at that lower frequency to contain some proportion of the higher

frequency (i.e. the input signal of amplitude A at frequency f2, corrupts the output at fi)[98].

41

The influence of cross-talk on the Zrs obtained using N T M signals has been confirmed by

Suki and Bates [101]. Using a nonlinear viscoelastic model, the authors demonstrated that

under conditions of increased ventilation amplitude and distortion the nonlinear cross-talk

increased such that it could no longer be regarded as negligible [101]. Furthermore, these

distortions will not be reflected in the y2 [98]. Non-sum non-difference pseudorandom

signals ( N S N D ) were recently introduced to minimize non-linear cross-talk [98]. However,

the authors demonstrated that if the non-linearities exceeded that of a second order system,

then it was necessary to use a 4th order N S N D signal (NSND-4) to completely eliminate

non-linear cross-talk from the impedance spectra estimated from a simulation study. The

investigators repeated this experiment in a single anaesthetized, tracheotomized and

paralyzed dog showing that the N S N D - 4 signal produced smooth estimates of Ers and Rrs

at all frequencies, while a N T M signal produced an impedance with smooth estimates of Rrs

and biased Ers estimates, whereas the composite signal produced Rrs and Ers values that

were clearly biased, particularly at low frequencies [98]. A further advantage of N S N D

signals is the reduced number of frequency components within the signal, allowing higher

energies to be assigned to the individual components within the signal. The disadvantage

of N S N D signals is the reduced resolution of the Zrs resulting from the smaller number of

frequency components and the longer recording times required for data acquisition. Indeed,

while N S N D signals may be appropriate for animal or human adults studies within an I C U

setting, the longer recording times required may preclude their use in spontaneously

breathing infants.

Inherent in any technique are assumptions. The forced oscillation technique assumes that

the measuring system and the mechanical properties of the respiratory system are linear

during the measurement. This implies that Zrs is independent of the amplitude of the

oscillatory signal. In intubated rabbits Peslin et al. showed that the application of small

amplitude pressure oscillations ( 2 - 4 c m H 2 0 ) at 10, 20 and 30 H z did not influence Rrs

or Ers. In addition the investigators demonstrated that the oscillatory signal did not

influence the lung response to histamine [100]. A second assumption is that Zrs values at

individual frequency components are equal to those obtained via single sinusoidal

oscillations. Desager et al. [102] showed matching Zrs measurements made using either

pseudorandom or sinusoidal signals at 12 and 32 Hz. A final assumption is that measured

42

flows and/or pressures are representative of the actual flows and pressures. If the upper

airway acts as a parallel shunt impedance then Zrs may be underestimated. Several studies

have sought to address this issue [21,103,104] and a detailed explanation is given below

(Section 1.3.1.1)

Although the work from animal studies can provide a rich source of technical and

methodological information, for ethical reasons a number of the techniques are unable to

be used in human subjects (e.g. alveolar capsules). As such the remainder of this section will

focus on those methods that have been successfully adapted for human use.

1.3.1 The forced oscillation technique in older children and adults

The impedance of the respiratory system can be determined in two ways, either using input

impedance (Zin) or transfer impedance (Ztr) systems. If the input and output signals are

applied and measured only at the airway opening, then the respiratory system can be

represented as a single node system (Zin). A two node system considers the airway opening

and body surface as separate nodes. In this case the forcing signal is applied at one node and

the resulting flows and pressures measured at the second node (Ztr) (Figure 1.2). With Zin

the oscillations are applied at the airway opening and the impedance determined from the

relationship of the pressure and flow measured at the airway opening (Zin = Pao/V). To

obtain the Ztr, oscillations are applied to the body surface and the impedance calculated

from the pressure surrounding the chest wall (Pbs) and airway opening flows (Ztr = Pbs/V).

In either case, the impedance spectra are plotted versus frequency and thus the frequency

characteristics of the respiratory system and its components (lung, chest wall, upper airways

and cheeks) can be determined. The advantages of the F O T are that the technique provides

a non-invasive way to measure the respiratory system, and that the ability to apply either

a single frequency or band of frequencies can provide useful information about the

pulmonary parenchyma. The disadvantages are that information is provided about all of the

respiratory structures, of which only the lung and chest wall are of primary interest and that

interpretation of the information is frequency dependent [94]. The use of selected

frequencies can overcome some problems as can the selective use of a model of the

respiratory system to partition the respiratory structures allowing those of interest to be

43

further studied. The advantages and disadvantages of the individual methods for

determining Zrs are outlined below.

/. 3.1.1 Input impedance

Initial studies measuring the input impedance of the respiratory system recorded data above

3 H z to avoid interference from the spontaneous breathing of the subject. Technical

advances have allowed the frequency range to be increased with current studies starting at

or around 2 Hz. These studies can provide general information about the mechanical

properties of the respiratory system, but are unable to characterize the frequency-dependent

lung tissues. Zin can be represented as a simple series model comprised of resistance,

inertance and compliance in these situations. Investigators have shown that by extending

the frequency at which Zin was determined out beyond 160 H z the first acoustic anti-

resonance of the airway system can be determined [105]. This extension of the frequency

being applied to the respiratory system requires a model accounting for tissue compliance

(Ct) and gas compression compliance (Cg) as suggested by DuBois [29]. This then allows

separate estimates of Raw, Rt and C g (see Figure 1.2) [106].

Only by extending the frequency range below the spontaneous breathing frequency can

information about the pulmonary parenchyma be obtained [3,16,96,107-109]. Recently, a

number of investigators have determined Zrs using frequencies below 2 H z [7,17,76,96,97].

Hantos et al. [96] investigated the frequency dependence of the human respiratory system,

lung and chest wall. In this study five adult subjects had their Zin measured in the 0.25-5

H z frequency range, with an esophageal catheter in place. The subject was required to

exhale and maintain an open glottis for 32s, enabling the characteristics of Zrs, Zl and Z w

to be determined. A negative frequency dependence was seen at low frequencies (< 2 Hz)

in both the resistance and the reactance for the respiratory system, lungs and chest wall

demonstrating that the respiratory system could be partitioned into lung and chest wall

components in healthy patients. This technique is limited by the effort and training required

to maintain an open glottis and would be unsuitable for younger children and infants.

Lutchen et al. [97] have successfully applied a waveform comprised of an oscillatory signal

superimposed on a tidal volume waveform. Termed the optimum ventilator waveform

technique ( O V W ) , it allows the measurement of low-frequency impedance data without the

44

need for an apnoeic period. This technique involves training the subjects to sit quietly while

relaxing to allow the application of the waveform. Lutchen et al. [97] have shown the

technique can reliably measure low-frequency input impedance spectra in healthy subjects

and in a single asthmatic showed the reversal of bronchoconstriction with the use of

bronchodilators. This technique requires less training than that required by the method of

Hantos et al. [96], but may be more suited to adults than to younger children, and would

be unsuitable in infants.

The application of the oscillatory signal at the airway opening may induce motion of the

upper airway walls and cheeks. The movement is responsible for some loss of measured

flow, leading to an underestimation of the respiratory impedance. Initial efforts to account

for this error involved the measurement of the upper airway wall impedance during Valsalva

manoeuvres [110], or by the use of a head plethysmograph to directly evaluate airway wall

motion during the Zin measurements [111]. These initial efforts have been improved upon

by the introduction of the head generator technique [21]. This technique involves applying

the oscillatory signal around the head rather than the mouth and thus minimizes the

transmural pressure across the extrathoracic airway walls during the respiratory impedance

measurement [21]. The authors found that the technique was almost as accurate as the more

cumbersome head plethysmographic technique and much simpler to implement. Since the

initial introduction of the head generator technique (HGT), several authors have made

direct comparisons between the standard input impedance technique and the head generator

technique in adults [103,104] and in children [104,112,113]. Cauberghs etal [104] found

no significant difference in Rrs between the H G T and standard techniques (cheeks

supported) in healthy adults, but the values of Rrs generated with the H G T were

significantly higher than that of the standard technique (cheeks unsupported), indicating a

mimmization of the upper airway shunt using the H G T . However, in adults with moderate

airway obstruction, the Rrs determined by the H G T was significantly higher than the

standard technique with a reduced frequency dependence. In addition, the resonant

frequency was significantly lower. In contrast, patients with severe airway obstruction

exhibited a negative frequency dependence with both techniques [104]. The authors

suggested that while the H G T did not completely suppress the upper airway shunt, it did

reduce the error when compared to the standard technique. These results have been

45

confirmed in a subsequent study, also in adults, that applied a frequency spectra of 8-256

H z and demonstrated that the H G T produced improved estimates of Zrs when compared

to the standard input impedance technique [114]. Habib and Jackson [103], using a

frequency spectra of 4-256 H z in 5 healthy adults measured Zrs using the standard, head

generator and water-filled head generator techniques to determine the input impedance (Zst,

Zhg and Zwa, respectively). The authors assumed that a water-filled head chamber would

eliminate upper airway wall movement due to the ̂ compressibility of water. Below 20 Hz,

the H G T provided nearly shunt-free estimates of Zrs, while above 32 H z the standard

cheek-supported technique resulted in a more accurate assessment of Zrs. Between these

frequencies the authors concluded that as both Zhg and Zst were different from Z w a it was

unclear which technique would provide the best shunt-free estimate of Zrs [103].

The error associated with the standard technique depends on the Zuaw/Zrs ratio, whereas

with the H G T the error is related to the ratio of impedance of the pneumotach (Zp) and

Zuaw (Zp/Zuaw). Marchal et al. [112] suggested that the H G T would provide better

estimates of Zrs than the standard technique with careful pneumotach matching and it was

subsequently shown in a group of normal children (3.5-13 years) that this was case. In a

further study, it was suggested that the H G T had a greater degree of sensitivity to detect

changes in lung function due to constrictor or bronchodilator agents in children (2.5-13

years) with a history of asthma [113].

1.3.1.2 Transfer impedance

To measure the Ztr of the human respiratory system, an oscillatory signal is applied to the

thorax of the subject while enclosed in a chamber and the resultant flow at the mouth

measured [94,115,116]. The description of the Ztr requires a more complex model than the

resistance-inertance-compliance (R-I-C) model (see Figure 1.1) used for input impedance

measurements. Ztr is usually analyzed using the six element model described by DuBois

[29] (Figure 1.2).

If the Cg is known, then the mechanics of the respiratory tissues can be determined. The

frequency range over which Ztr is measured greatly influences the reliability of the airway-

tissue separation [117]. Recent studies have suggested that the frequency range needs to

46

be extended at least to 50-64 H z [6,118]. If frequencies of 4-64 H z are used, then Ztr can

provide several advantages over Zin [117], as the latter cannot provide specific information

relating to the airways and tissues (see Section 1.3.1.1). In addition, measurements of Ztr

should be less sensitive to the errors associated with upper airway shunting [116]. However,

the ability of Ztr to accurately determine the behaviour of the pulmonary parenchyma has

been recently questioned by Lutchen et al [116]. The authors found that in a group of

adults with various forms of lung disease that the bronchodilator-induced changes in Rt

were uncorrelated with changes in spirometric values. However, when compared to

normals, the authors found that this indicator (Rt) was sensitive to disease [116]. These

results question the ability of Ztr to provide a reliable clinical tool to assess changes in lung

tissue mechanics, leading to the conclusion that Ztr m a y only be suitable as a non-invasive

method for assessing changes in airway function [94].

1.3.2 The use of the forced oscillation technique in infants

The translation of the Zin and Ztr techniques to an infant population has provided a number

of obstacles, the largest of which is that the population is inherently uncooperative. While

a few studies have been carried out in unsedated infants [119], the majority of studies have

been carried out in infants either within the intensive care setting or under sedation.

1.3.2.1 Input impedance

The two forms of input impedance described above (standard and H G T ) have both been

adapted for use in infants. The studies within each of these areas can be split into the signal

frequencies that have been applied with the majority of studies using the standard technique

carried out in the 4-64 H z range [102,120]. Desager et al [102] adapted the standard input

impedance technique for use in infants and demonstrated that reliable estimates of Zrs could

be obtained between 16 and 52 Hz. The authors evaluated the linearity of the measurement

apparatus and of the infant respiratory system and found them to be linear within the range

of measurement. In addition inter- and intrasubject variabilities were acceptable ( 6 % and

5.7%, respectively). The authors also showed that after spontaneous turning of the head,

and subsequent repositioning that there was no significant difference in impedance values.

In a further study, the nasal impedance was evaluated [120] from 4 - 52 Hz, showing R n

contributed up to 4 8 % of Rrs at 24 H z . Jackson et al [106] have investigated the high-

47

frequency behaviour of the respiratory impedance in 9 healthy infants, characterizing the

input impedance from 4 - 1 6 0 Hz. The authors found that below 20 H z the coherence

functions were poor and hence only applied the DuBois model between 20 and 160 Hz. The

authors found that using the DuBois 6 element model and a subsequent 10 element model,

while the resultant respiratory parameters were unduly influenced by the shunt impedance

of the face mask, Rrs estimates were related to airway resistance [106]. Frey et al. [121]

determined baseline high-frequency Zrs measurements in 13 wheezy infants and assessed

the changes due to methacholine constriction in 9 infants. The authors found that the first

anti-resonance could be reliably determined in all infants and that it increased significantly

following constriction and concluded that these measurements may offer insight into

changes in airway wall compliance during constriction. The mechanical behaviour of the

respiratory tissues are frequency-dependent and dominate Rrs and Xrs below the

spontaneous breathing frequency. It would be useful therefore to characterize the behaviour

of the pulmonary parenchyma below 2 - 4 Hz.

Marchal et al. [122] have adapted the HGT to determine Zrs in infants. In this investigation,

the authors compared the resistance and compliance obtained using the H G T and the

standard technique. The authors found that if nasal impedance is accounted for, the error

in the resistance determined by the H G T was small compared to the standard technique.

Estimates of compliance were not significantly different. Normally the standard technique

will underestimate Rrs, while the H G T will overestimate Rrs. In nasal breathers (e.g.

infants) this overestimation of Rrs may be exaggerated [122].

Clinical studies examining changes in low-frequency Zrs with inflation pressure [76] and

following bronchodilators in healthy and wheezy infants have been conducted [123]. Zrs

has also been determined in infants within an I C U setting [124-126].

1.3.2.2 Transfer impedance

The first use of Ztr in infants was carried out by Wohl et al. [127] in 1969. The authors

applied a forcing signal between 3-7 H z around the body and demonstrated inspiratory

resistance to be lower than expiratory resistance. Marchal et al. [128] measured Ztr and

thoracic gas volume (TGV) in 22 asymptomatic infants. In 12 infants the impedance data

48

was fitted to the DuBois six element model, providing estimates of specific Raw, Rt, law

and Ct. Using simulation studies, the authors showed that the poor fit in the remaining

infants could be attributed to an upper airway shunt during nasal breathing [128]. More

recent work by Jackson and co-workers [129], however has shown that the frequency range

used to obtain Ztr was crucial in the DuBois model was to be utilized to estimate the airway

and respiratory tissue properties in infants. The investigators demonstrated than Ztr may

be measured over a frequency range of 4-140 Hz, and that this frequency range allowed

reliable estimates of R a w and Rti, using the DuBois model, in spontaneously breathing

infants.

1.5 The low-frequency forced oscillation technique in infants

In adults and older children the detailed investigation of spirometric techniques and large

study numbers have enabled the development of predictive equations allowing disease

severity to be quantified. Numerous techniques have been developed to ascertain lung

function in younger children and infants, but the current literature demonstrates that most

methods of determining lung function are unable to provide separate estimates of the airway

and lung tissue mechanical properties. One technique able to determine airway and

mechanical properties simultaneously is the low-frequency forced oscillation technique [7].

This technique applies small-amplitude pseudorandom pressure oscillations to the airway

opening and input impedance is determined and the constant-phase model subsequently

fitted [4]. Previous studies have characterized the volume dependence of the airway and

lung tissue mechanics [76], and assessed the effect of bronchodilators on healthy and

wheezy infants [123]. These studies have not addressed such issues as the effect of the nasal

impedance on the input impedance of the respiratory system or the contribution of the chest

wall to Zrs. Factors affecting measurement conditions need to be standardized, including

the most appropriate signal type (EvI vs. N T M or N S N D ) , measurement time and signal

analysis, head position and neck extension or flexion, the ability of the low-frequency F O T

to provide diagnostic information and changes associated with normal growth in infants.

1.6 Aims

The aims of this thesis were to:

49

1) establish the ability of the low-frequency F O T to non-invasively partition the respiratory

system into lung and chest wall components and to use a linear model of the lung to

characterize the airway and parenchymal mechanics. This was initially carried out in an

animal model before being translated into children and infants.

2) determine the type of oscillatory signal required (IM vs. NIM) in terms of model fitting

criteria and the success rate of obtaining technically acceptable Zrs data.

3) characterize the contribution of the nasal and chest wall impedances on the respiratory

system impedance in infants.

4) measure airway and lung tissue mechanics in healthy infants over the first two years of

life and hence model the relationship between mechanical behaviour of the respiratory

airways and tissues and length using the low-frequency FOT.

5) compare the airway and pulmonary tissue mechanics of wheezy infants with those of

normal infants and hence determine the ̂ criminatory power of the technique.

6) determine the effectiveness of the low-frequency FOT to detect methacholine

responsiveness in infants.

50

51

CHAPTER 2: METHODS AND MATERIALS

2.1 Introduction

This chapter outlines the equipment, techniques and protocols used in the remainder of this

thesis. There are a number of adaptations of the forced oscillation technique (FOT), many

of which have been described in the review of the literature. All of the work within this

thesis has been determined by obtaining the input impedance (Zin). Either a wave-tube or

a pneumotachograph have been used to measure Zin. The first allows Zin to be obtained

without the need to measure flow, the use of a wave-tube is more suitable in small animal

models where the respiratory flows are low. The second measures the flow and pressure at

the airway opening and hence Zin. In addition, two set-ups, utilizing pneumotachographs,

have been used, the first in spontaneously breathing infants and the second in anaesthetized,

paralyzed and mechanically ventilated patients. The mechanical properties of the airways

and tissues have been characterized using the constant-phase model of the lung, the

background, limitations and advantages of the model are outlined.

2.2 Measurement techniques for determining respiratory

impedance

2.2.1 Wave-tube device

The measurement setup used to collect the low-frequency impedance data was an

adaptation of that used by Petak et al [5] and is shown in Figure 2.1.

2.2.1.1 Validity of the wave-tube device

The wave tube device does not measure flow and as such the input impedance is determined

from the transfer function of the loudspeaker pressure and the airway opening pressure. In

an open system the transfer function will be negligible, while in a closed system the transfer

function will approach unity. As described in further detail in Section 2.3 the forcing signal

is constructed so as to provide the best possible signal to noise ratio. The physical

characteristics of the wave tube are determined to maximise the sensitivity of the wave-tube

to detect changes in the load impedance while minimising the possible influences of

artefacts. The diameter and length of the wave tube are chosen based on the following:

52

Diameter is adjusted to the tracheal or cannula ID. This is a strong constraint as resistance

is extremely sensitive to the diameter (~dA(-4)). Length is determined by the expected load

impedance and the wavelength. For example if L is too short then P2/P1 approaches 1 at

the lowest frequency, and the computation becomes too sensitive to calibration and

temporal stability. Alternatively, if L is too long, P2 becomes very small and may be subject

to quantisation and other noises. A compromise in L has to be found. The other determinant

is the L vs. wavelength. L is kept below the quarter of the wavelength in order to avoid

standing waves and reflections. Here the highest-frequency component matters. For

example for a 25 H z maximum frequency the wavelength is 330 m/s (sound velocity in air)

over 25/s = 13.2 m, hence the quarter wavelength is 3.3 m. The 114 cm wave-tube used is

well below this upper limit.

2.2.1.2 Measurement apparatus

The measurement circuit consisted of the animal, the ventilator and the measurement

system, the components were joined at a 3-way tap that could be switched between the

measurement system and ventilator. The measurement system included the loudspeaker-in-

box: used to generate the oscillatory signals, a 114 cm long 2 m m ID polyethylene wave-

tube of known characteristics and 2 identical miniature pressure transducers (ICS model

33NA002D) to determine the lateral pressures at the loudspeaker (Pi) and the tracheal end

(P2) of the tubing. Esophageal pressure (Pes) was measured with respect to atmosphere to

separate Zrs into Z w and Zl via positioning a 3-french single-sensor catheter-tip

micromanometer ( M T C 3F Drager Medical Electronics, Best, Netherlands) into the

oesophagus. The pressure drop across the chest wall was monitored during mechanical

ventilation, and the catheter was positioned to obtain a smooth respiratory curve with

minimal cardiac noise. The position of the catheter was finalized by performing the 'positive

pressure' occlusion test described by Lanteri etal [130]. The animal was placed inside a

plethysmograph and the loudspeaker connected (Figure 2.1), the airway was occluded at

the 3-way tap and the oscillatory signal applied to the body surface. The resultant changes

in P2 and Pes were examined to confirm catheter placement. To obtain impedance

measurements the E T tube was switched to the loudspeaker-in-box system at end-

expiration. The loudspeaker generated a small-amplitude pseudorandom signal containing

15 non-integer multiple (NTM) components in the frequency range of 0.5-21 H z through

53

the wave-tube.

P, P-1 ' 2 Figure 2.1 Set-up for the wave-tube technique. The dotted lines show the setup for the positive-pressure

occlusion test.

The Pi, P2, and Pes signals were low-pass filtered, and digitized by an analog-digital board

of an IBM-compatible computer at a sampling rate of 128 Hz. The pressure transfer

functions Pi/P2 and Pes/P2 were computed by fast Fourier transformations (FFT) from the

6-s long recordings, by using 4 s time window and 9 5 % overlapping. Zrs was calculated as

the load impedance of the wave-tube [5]:

Zrs = Zo sinh(vL)/{(Pi /P2 -cosh(yL)} -(27)

where: L is the length of the tube. The characteristic impedance (Zo) and the complex

propagation wave number (y) were determined by the geometrical data and the material

constants of the tube and air.

Since tracheal flow was not measured, Z w was calculated assuming no flow loss to airway

wall or alveolar gas compression as Z w = Zrs (Pes/P2) and Zl was obtained by subtraction

(Zl = Zrs-Zw). The load impedance of the E T tube and the connecting tubing was also

determined.

54

2.2.2 Pneumotachograph device

This setup involves the measurement of central flow and pressures (transrespiratory and/or

transpulmonary) to determine Zrs, Zl and Zw. Zin was obtained in spontaneously breathing

or mechanically ventilated patients and the methods used are detailed below.

iii

Figure 2.2: Setup for the technique used for spontaneously breathing infants. During forced inspirations the

valve between the loudspeaker and reference chamber was open, allowing equalization between the loudspeaker,

the infant and the reference chamber. Following the final inflation, the balloon valve and valve was closed, hence

occluding the airway and isolating the loudspeaker from the reference chamber.

2.2.2.1 Spontaneously breathing patients

Respiratory input impedance spectra (Zrs) were measured by using the low-frequency F O T

during an apnoeic pause induced by the Hering Breuer reflex as described by Sly et al. [7].

The measurement setup is shown in Figure 2.2 and included the loudspeaker-in-box system,

a reference chamber and the computer controlled pump and balloon valve, the two

components were connected to a soft-rimmed face mask using a T-shaped connector. The

oscillatory forcing signal was generated by a loudspeaker-in-box system and was introduced

to the infant, via a soft-rimmed face mask fitted to ensure a leak-free seal around the nose

55

and mouth. The Hering Breuer reflex was induced by introducing inspiratory flows

generated by a computer controlled pump, in series with the balloon valve. Prior to

obtaining the oscillatory measurements, three deep forced inspirations were applied (Figure

2.3): the balloon valve was opened and the inspiratory flow was maintained by a pump

(Inflate-all; Coleman Co. Inc., Wicket, K S ) until the transrespiratory pressure (Pao) reached

20 c m H 2 0 , each inflation was followed by a pause of 0.75 s to allow passive expiration,

before the successive forced inspiration was applied. At the end of the third inflation, the

balloon valve was closed, occluding the airway at a transrespiratory pressure of 20 c m H 2 0

and hence inducing the. If three successive inflations were insufficient to induce the Hering

Breuer reflex the number of inflations was increased to a maximum of eight inflations. In

the resulting short apnoeic period, the valve between the loudspeaker box and the reference

box was closed and the low-frequency pseudorandom signal was driven into the infant's

respiratory system by the loudspeaker. The two chambers of the loudspeaker box,

pressurized during the inflation maneuvers to keep the membrane in the midposition, were

connected to each other and to the reference box through a T-tubing. This allowed for rapid

equilibration of static pressures, but did not attenuate the amplitude delivered at oscillation

frequencies. Measurements were 6s in length and the Pao recording examined for leak or

respiratory efforts, with corrupted recordings being excluded.

The oscillatory signal contained 16 frequency components in the 0.5-20 Hz range. The

relative amplitudes of the lower frequency components were elevated to maintain sufficient

signal-to-noise ratios at all frequencies and the signal optimized to keep the peak-to-peak

amplitude of the signal to a minimum. V was measured with a screen pneumotachograph

(28 m m ID) connected to an ICS 3 3 N A 0 0 2 D differential pressure transducer (ICSensors

Inc., Milpitas, CA). A n identical pressure transducer was used to sense respiratory pressure

(Prs) as the pressure difference between the mask and the reference box. The latter was

used to eliminate the 20 c m H 2 0 static pressure and hence, increase the resolution. Nasal

pressure (Pn) (if measured: See Chapter 6) was determined using a solid state miniature

tipped catheter ( M T C 5F; Drager, Best Netherlands) placed 4-5 cm into the nasal cavity,

slightly above the nasopharynx. The signals of Pn, Prs and V were low-pass filtered at 25

Hz, sampled at 128 H z with a 12-bit analog-digital converter and stored on an IBM-

compatible computer for later analysis.

56

O 20 CN

0 5 10 15

Time (s) Figure 2.3: Time domain trace of mouth pressure depicting the 3 inflations and the airway occlusion.

20

Three to ten measurements of the uncorrected input impedance (Z*) were collected in each

infant. Following each measurement the mask was lifted and the infant allowed to breathe

spontaneously. Following each series of measurements the impedance of the apparatus dead

space (Zds) was determined. Z* and Zds were computed from the cross-power spectra

between the measured V and Prs signals and the stored driving signal using the L U N G

software (SZOTE University, Szeged, Hungary [96]. The spectra were obtained by fast

Fourier transformation (FFT) from 3 s time windows and 9 5 % overlapping. Zds was

considered as a lumped shunt impedance in parallel with the respiratory system, and the

corrected respiratory impedance (Zrs) was calculated as Zrs = (ZdsZrs*) / (Zds - Zrs*). A

linear model of the respiratory system [26] was then fitted to the corrected individual

respiratory impedance spectra, in the 0.5-15 H z range (see Section 2.4). The same

procedure was followed when determining the lower respiratory system impedance (Zlrs).

57

2.2.2.2 Mechanically ventilated patients

The measurement set-up used to collect Zrs data on spontaneously breathing infants [7] was

modified to measure impedance in the intubated children (Figure 2.4). Hand operated valves

were used to switch the ET-tube from the respirator to the measurement apparatus as

follows. Valves A and C were open and B closed during mechanical ventilation. Preceding

oscillatory measurements the lungs were inflated to a pressure of about 30 c m H 2 0 to

standardize the volume history. The valve B was opened for a few ventilatory cycles before

data collection to equilibrate the pressures between the children and the measurement set

up. The valve A was then closed at end-expiration, and the communication between the

loudspeaker chamber and the reference box was suspended by closing valve C. In the

resulting apnoeic period, small-amplitude (<2 c m H 2 0 peak-to-peak pressures) pseudo

random pressure excitations were introduced into the trachea. The forcing signal contained

thirty integer-multiple components of the fundamental frequency 0.4 Hz, resulting in a

frequency range of 0.4-12 Hz. The energies of the low-frequency components were

enhanced to provide a sufficient signal-to-noise ratio over the whole frequency range. The

phases of the components were optimized to give a minimal peak-to-peak pressure

excursion at the airway opening. 4-6 recordings were collected in intact-chest condition.

Periods of at least 2 min of mechanical ventilation were interposed between the successive

measurements.

Tracheal flow (V) was measured with a 28-mm-id screen pneumotachograph connected

to a differential pressure transducer (ICS 33NA002D). To avoid the inclusion of the E T -

tube impedance into the measurements, intratracheal pressure (Ptr) was measured with an

identical pressure transducer through a 1.5-mm-OD polyethylene catheter positioned 1.5-2

c m distal to the ET-tube. The reference port of this transducer was connected to a reference

box to eliminate the D C offset. Airway opening pressure (Pao) with reference to the

atmosphere was also detected with a third ICS sensor to monitor absolute tracheal

pressures. To separate Zrs into Zl and chest wall (Zw) components in the intact-chest, the

esophageal pressure (Pes) with reference to the atmosphere was measured by introducing

a 3 french catheter-tip micro-manometer ( M T C 3F Drager Medical Electronics, Best, The

Netherlands) into the lower third of the esophagus.

58

loudspeaker

Figure 2.4: Pneumotach setup used in mechanicahy ventilated patients. During normal ventilation valves A and

C are open, with valve B clamped shut At end expiration valves A and C were shut and valve B opened allowing

the forcing signal to be applied. Tracheal pressure (Ptr) was determined distal to the ETT.

The pressure drop across the chest wall was monitored during mechanical ventilation, and

the catheter was positioned to obtain a smooth respiratory curve with minimal cardiac

noise. The electrical signals of the transducers were low-pass filtered at 25 Hz, and digitized

at a rate of 256 H z by an analog-to-digital board of an IBM-compatible personal computer.

The correction spectra for each transducer pair were determined by exposing the sensors

together with their connecting tubing to the same oscillatory pressure fluctuation. The

differences in the frequency response of the pressure transducers and their tubing measuring

Ptr, Pes, and V were recorded, and the impedances were corrected accordingly.

The mechanical impedance of the respiratory system was calculated from the Ptr and V

signals (Zrs = Ptr/V1). The chest wall impedance was computed directly from the Pes and

V signals (Zw = Pes/V), while Zl spectra were obtained by subtraction (Zl = Zrs-Zw). The

59

impedance spectra were calculated by fast Fourier transformation from the 7.5 s long

recordings using 2.5 s time window and 9 5 % overlapping.

2.3 Forcing signals

The selection of forcing signal is crucial to ensure a maximization of the signal to noise ratio

(S/N). Studies conducted in mechanically ventilated, paralyzed patients have sufficiently

long apnoeic periods for the longer N T M or N S N D signals to be used. Within spontaneously

breathing infants the apnoeic period is elicited using the Hering-Breuer reflex. Apnoeic

periods longer than 6 seconds cannot always be obtained, hence restricting the signal type

that may be utilized. In order to determine the optimum forcing signal to use a pilot study

was conducted in 6 spontaneously breathing infants and children. In each patient 2 signals

were applied in random order. Signal 1 was an I M signal and contained 16 frequency

components, with a fundamental frequency of 0.5 H z (T=2s and T = signal period), in the

0.5-20 H z range. Signal 2 contained 23 non-integer multiple (NTM) components, with a

fundamental frequency of 0.25 H z (T=4s), in the frequency range of 0.5-21 H z (Table 2.1).

4-6 Zrs measurements were obtained in each patient (as described above). If the

measurement period is greater than or equal to 3 times the signal period, then a fast Fourier

transform (FFT), with a window of twice that of the signal period may be used to convert

the time-domain signal into the frequency domain (double FFT). The resolution of the

resulting frequency domain signal will be 1/2T (e.g. if T = 2, then the resolution will be 0.25

' Hz). Measurement times less than 3T only allow for the use of a FFT, with a window of T

seconds and hence a resolution of 1/T (single FFT). The better the resolution the lower the

chance that noise or artifact will be included in the frequency estimate (i.e. improved S/N

ratio). In addition FFTs were 9 5 % overlapped with the successive FFT. Overlapping

describes the overlap between one Fourier transform window and the next. This indicates

that the second transform overlaps the first by 95%. The effect of large overlapping is to

improve the signal to noise ratio of the data, as noise will be averaged to zero, while the

data will remain unchanged. Signal 1 was analyzed in the following way. If no breath

occurred during the measurement, then a double FFT was used. If a breath occurred during

the measurement between t =3-6s, then the portion of signal containing the breath was

excluded and a single FFT performed on the remaining signal. If the breath occurred prior

to 3 seconds then the measurement was excluded.

60

Signal 2 was analyzed using a single FFT. Individual Zrs spectra were fitted to the constant-

phase model of the lung (see Section 2.4). In addition the success rate was calculated for

each infant. Paired students t-tests were applied to the mean respiratory parameters, the

variability (SD), the fitting error (F%) and the success rate.

Measurements for each signal were obtained in all children. There was no significant

difference in any respiratory parameter or the variability in any parameter. The fitting error

obtained from the constant phase model when using signal 1 (7.6±0.6%) was significantly

lower (p<0.01) than that obtained when using signal 2 (8.6±0.5%). This lower fitting error

may be due to the use of the double F F T possible with the shorter I M signal. The shorter

measurement time also resulted in a non-significant increase in success rate when signal 1

was used (79±3% and 65±6%, for signal 1 and 2, respectively)

The use of NTM and NSND signals ensure decreased artifact due to cross-talk and non-

linearities [98], however the increased measurement times required introduce practical

difficulties. In adults and ventilated patients the potential to induce a pause in breathing, of

sufficient length to obtain the measurement times required for N T M or N S N D signals is

increased. In uncooperative infants the pause in breathing must be induced by invoking the

Hering-Breuer reflex. As this reflex decreases with age, the ability to obtain the minimum

pause needed for a signal, such as Signal 2 (above) is decreased. In infants with an increased

respiratory drive, such as during an acute exacerbation of wheeze, or during a methacholine

challenge test the length of the pause in breathing will be reduced. The use of the shorter

EvI signal will allow for a higher success rate in all patients groups, the additional use of a

double F F T ensure high S/N ratios. The subsequent studies contained within this thesis use

I M signals.

61

Table 2.1: Frequency components of I M and N I M signals

IM(Hz)

0.5

1.0

1.5

2.0

3.0

4.0

5.0

6.0

7.5

9.0

10.5

12.0

14.0

16.0

18.0

20.0

N I M (Hz)

0.50

0.75

1.25

1.75

2.75

3.25

4.25

4.75

5.75

7.25

7.75

9.25

10.25

10.75

11.75

13.25

14.75

15.25

16.75

17.75

18.25

19.75

20.75

2.4 Constant-phase model

Impedance spectra were evaluated in terms of model parameters. The model contained a

frequency independent resistance (R) and inertance (I) connected in series with a constant-

phase tissue compartment [26] representing tissue damping (G) and elastance (H):

Z = R + jcol + (G - jH)/oa -(28)

where j is the imaginary unit, © is the angular frequency, and a is expressed as a = 2/n

arctan (H/G).

62

Figure 2.5 shows a schematic representation of the mechanical behaviour of the airways and

the respiratory tissues and their contribution to the frequency dependent behaviour of the

respiratory system as described by the constant-phase model. The inverse frequency

dependence of the respiratory tissues can be seen at low frequencies (< 2 Hz). The

resistance of the airways are frequency independent and contribute a majority of the

respiratory resistance at higher frequencies (>10 Hz), while the inertance of the airways

make a negligible contribution to the respiratory system reactance (Xrs) at low frequencies

and dominate Xrs above 10 Hz.

When the model is fitted to the Zrs, Zl or Zw spectra, the tissue parameters characterize the

damping and elastic properties of the corresponding tissue compartment. For the pulmonary

impedances, the R and I reflect primarily the airway resistance (Raw) and inertance (law).

In the case of the Z w , the R and I represent the Newtonian resistive and inertive

components of the chest wall. W h e n the model is fitted to Zrs spectra, the estimated R and

I contain R a w and law and the resistive and inertive contribution of the chest wall. The

model parameters were obtained by using a fitting criteria by minimizing the following

objective function:

F = f(I/m)j^f\ 2(ayi)-Z'(coi)V\ Kto,)?]?2

where F is a measure of the 'goodness-of-fit' of the model.

30 -,

15 -

0 -

63

Frequency (Hz)

Figure 2.5: Schematic representation of the airway (dotted line) and tissue (dashed line) contributions to the

frequency dependent behavior of the respiratory system (solid line).

64

2.5 Raised volume rapid thoracic compression technique

Forced expiratory data were obtained over an extended volume range using the technique

described by Hayden et al [131]. This technique uses a pump (Inflate-all; Coleman Co. Inc.,

Wicket, K S ) to raise the infant's lung volume above the tidal range. The pump draws room

air through a pneumotachograph (No. 3719; Hans Rudolf Inc, Kansas City, M O ) and into

the infant via a soft-rimmed face mask (King System, Noblesville, IN). Flow (V) and airway

opening pressure (Pao) were measured, amplified (SC14C; RHT-INFODAT, Montreal, PQ,

Canada) and recorded on a computer using B R A T L A B data acquisition and analysis

software (RHT-INFODAT). A series of computer controlled solenoid valves were used to

inflate the infant's lungs 3 times to a transrespiratory pressure of 20 cmH 2 0 with passive

deflations between each inflation. Using a jacket connected to a positive pressure reservoir,

a compression force was applied to the thorax and abdomen after the third and final

inflation. The compression force was standardized to transmit a pressure of 20 c m H 2 0 to

the airway at end-inflation resulting in an airway opening pressure of 40 cmH 2 0 [131]

(Figure 2.6).

Forced expiratory flow was measured and volume calculated. FEV0.5, FEV0.75 and FEVi

were calculated from the forced expiratory volume-time curves. Three to ten technically

correct measurements were obtained from each infant.

65

pneumotachograph

400 kPa air supply

Hand operated valve

Pressure I

pressure guage

valve (2)

To jacket

Pressure reservoir

Figure 2.6: Schematic of R V R T C measurement set up. R o o m air is drawn through the pneumotachograph by

the pump via valve 4 and forced into the infant's lungs via valve 3. Valves 3, 4 and the balloon valve act in

concert to allow the forced inflations and passive deflations. The jacket is inflated from the pressure reservoir

via valve 2, following forced expiration the jacket empties through valve 1.

66

67

CHAPTER 3: REPEATED MEASUREMENTS OF AIRWAY AND LUNG TISSUE MECHANICS

IN RATS

3.1 Summary

For studies investigating the mechanisms underlying the development of allergic conditions

such as asthma, non-invasive methodologies for separating airway and parenchymal

mechanics in animal models are required. To develop such a method, seven Brown Norway

rats were studied on three occasions over a 14-day period. Following the third intact chest

baseline measurement inhaled methacholine was adniinistered. Once lung function returned

to the baseline level, a thoracotomy was performed to compare the lung mechanics in the

intact and open chest conditions. O n each occasion, the rats were anesthetized, paralyzed

and intubated. Small amplitude pseudorandom oscillations between 0.5-21 H z were applied

through a wave tube to obtain respiratory impedance (Zrs). Esophageal pressure was

measured to separate Zrs into pulmonary (Zl) and chest wall (Zw) components. A model

containing a frequency independent resistance and inertance, and a tissue component

including a tissue damping and elastance was fitted to Zrs, Zl, and Z w spectra. N o

statistically significant change was found in any Zrs, Zl, or Z w parameters between tests.

The number of animals required to show group changes in lung mechanics was significantly

less when measured non-invasively than when calculated from open chest parameters. In

conclusion, the current method can be used to separate airway and lung tissue mechanics

non-invasively over a series of tests and can detect pulmonary constrictor responses for the

airways and the parenchyma separately.

68

3.2 Introduction

Recent studies in both animals and humans have established that the pulmonary parenchyma

plays an important role in determining lung function and in determining the mechanical

response to various insults. Total lung resistance (Rl) comprises two components, the flow

resistance of the bronchial airway tree against the flow (Raw) and the resistance of the

pulmonary parenchyma (Rti) manifested as a viscoelastic pressure loss across the lung

tissues. Animal models have demonstrated that changes in lung function can be produced

by changes occurring predominantly in the airways, in the parenchyma, or in both

compartments.

The major problem limiting wider application of these findings to fundamental research

questions in animals and humans is the invasive nature of the methods used to date. Studies

performed in animals have required direct measurement of alveolar pressure using pressure

capsules glued to the lung surface, usually in open-chested animals, or in-vitro methods.

These methods preclude repeat studies in a single animal which would be desirable for many

of the studies likely to be required for the successful development of a realistic animal

model of asthma. Thus the development of a non-invasive method for partitioning lung

mechanics into airway and parenchymal components would represent a major advance.

Allowing the detailed study of lung function required to understand the pathogenesis of

lung diseases such as asthma.

Clinical longitudinal studies on the effect of aging [132,133] or to explore the long-term

development of a disease process [134] are frequently performed to follow changes in the

mechanical properties of the respiratory system. For the same purpose in animal studies,

independent groups are used and long-term changes in the lung function are detected from

unpaired comparisons. Nevertheless, due to the baseline variability and/or the apparent

scatter in the lung responsiveness, changes occurring long-term in the lung mechanics may

remain undetectable or require large number of animals in the independent groups.

The primary aim of the present study was to develop and test a modified forced oscillatory

technique that allows the partitioning of airway and parenchymal mechanics non-invasively,

and which would be suitable for performing long-term follow-up studies in rats. The rats'

69

low-frequency total respiratory impedance spectra (Zrs) was partitioned into components

representing the lungs (Zl) and the chest wall (Zw). Airway and parenchymal mechanics

were separated from Zl using a model containing an airway and a constant-phase [4] tissue

compartments. The baseline variability of the airway and lung tissue mechanical parameters

were determined in a two-week follow-up.

3.3 Methods

3.3.1 Animal preparations.

Seven male Brown-Norway rats were studied (275±17 g). The animals were housed in a

pathogen-free colony and allowed food and water ad libitum. Anaesthesia was induced with

an intraperitonial injection of ketamine (90 mg/kg) and rompun (5 m g /kg) mixture, and an

endotracheal tube (6-cm long, 2-mm ID) was inserted. The rats were then placed in the

supine position and mechanically ventilated (Model 683: Harvard Apparatus, South Natick,

M A ) with a frequency of 90 breaths/min and 9 ml/kg tidal volume. The tail vein was

cannulated for intravenous (iv) drug delivery by introducing a butterfly needle through a

small puncture on the skin, and pancuronium bromide (0.4 mg/kg) was administered to

induce muscle paralysis. Maintenance doses of the anaesthetic mixture and the pancuronium

bromide were administered iv every 40 minutes, or as needed. A n iv bolus of atropine (50

ug/kg) and neostigmine (100 Mg/kg) was used to advance recovery following the

measurements.

Following the completed non-invasive measurement protocol a two-step thoracotomy was

performed [135,136] to estimate the intact chest end-tidal transpulmonary pressure (Ptp),

which was then set as a positive end-expiratory pressure (PEEP) in the open chest

measurements. Briefly, the diaphragm was exposed from the abdomen side. The trachea

was then occluded, and the step change in the tracheal pressure was recorded resulting from

the bilateral cut of the parietal pleura. The magnitude of this step change (1-2 c m H 2 0 ) was

taken as the in situ end-expiratory pressure, and this was set to keep the same end-

expiratory lung volume in the open chest condition.

70

3.3.2 Measurement apparatus

See Section 2.1.1 for full details.

The measurement setup used to collect Zl data was described in detail previously [5].

Briefly, the E T tube was switched to a loudspeaker-in-box system at end-expiration. The

loudspeaker generated a small-amplitude pseudorandom signal through a polyethylene

wave-tube. Esophageal pressure (Pes) was measured with respect to atmosphere to separate

Zrs into Z w and Zl. Zrs was calculated as the load impedance of the wave-tube [5]. Z w

was calculated assuming no flow loss to airway wall or alveolar gas compression as Z w =

Zrs(Pes/Ptr) and Zl was obtained by subtraction (Zl = Zrs-Zw). The load impedance of the

E T tube and the connecting tubing was also determined.

3.3.3 Study protocol

Respiratory mechanics were measured on days 0, 7, and 14. O n each occasion, 6-8 data

epochs were collected, and the Zrs^ Z w , and Zl spectra were ensemble-averaged. O n day

14, after the baseline measurements were collected, aerosolized M c h (2 mg/ml) was

administered with a jet nebuliser (LC PLUS, Pari-Werk, G m b H , Germany) driven by 51/min

air attached to the inspiratory port of the ventilator for 60 seconds. Data collection was

started 1 min after completion of M c h challenge and successive recording were taken every

minute thereafter. Individual Zl and Z w curves were calculated and parameter values at the

peak response in Zl were reported.

Following the Mch challenge, lung mechanics were allowed to return to the baseline. The

chest was then opened, and PEEP was applied to maintain the end-expiratory lung volume

as close as possible to that occurring with the chest intact (see above). The pressure in the

loudspeaker box was adjusted to the P E E P level during these measurements to keep Ptp

unchanged. The open-chest pulmonary impedance spectra (Zlo) was determined as the load

impedance of the wave-tube.

One rat failed to recover following the second test, six rats completed the two-week follow-

up and form the basis of the present report.

71

3.3.4 Parameter estimation

See Section 2.2 for full details.

A linear model containing a frequency independent resistance (R) and inertance (I) and a

tissue damping (G) and elastance (H) of a constant-phase tissue compartment [4] was fitted

to the Zrs, Z w , Zl, and Zlo spectra by minimizing the absolute difference between the

measured and the modelled impedance data. Tissue hysteresivity (r|) [137] values were

calculated as G/H. Impedance data coinciding with the heart rate or its harmonics had poor

reproducibility and were excluded from the model fit.

3.3.5 Statistics

Scatters in the parameters were expressed in SE values. Repeated measures of one-way

analysis of variances with the Student-Neumann-Keuls multiple comparison procedure was

used to compare the parameter values in the different states of the study. Sample size

calculations were performed as follows. For the intact chest lung data, w e assumed that the

changes will be detected within an individual rat. Therefore, the minimum sample sizes were

determined by using the standard procedure of the paired t-test. In this case, the expected

variability of the changes in the intact chest Zl data was estimated from the 14 day follow-

up of the intact chest lung parameters. The assessment of the sample sizes for the open

chest protocol assumed two separate groups of animals. Accordingly, the standard

deviation values of the lung parameters were determined from the open chest lung

impedance data, and the minimum sample sizes were calculated from an unpaired t-test. In

both cases, changes of 1 0 % , 2 0 % , and 5 0 % were expected in Raw, G, and H. The power

of the tests was 0.8 and the coefficient of variation was kept constant. For sample size

calculation, the resistance of the E T tube and the connecting tubing ( R E T = 98 cmH20.s/L)

was subtracted from the R a w values.

3.4 Results

3.4.1 Non-invasive partitioning of Zrs

Representative Zrs, Zl, and Z w curves with the corresponding model fits are demonstrated

in Figure 3.1. All impedance curves exhibit similar frequency dependence. The low

72

variability seen across all frequencies, except those coinciding with the heart rate or its

harmonics, indicate highly reproducible impedance measurements. Zl contributes the

majority of the high frequency real part of Zrs suggesting that the lung dominates the

Newtonian resistive properties of the respiratory system. Z w comprises most of the increase

in the real part of Zrs at low-frequencies indicating that the chest wall has a major

contribution to G. In contrast, the low-frequency imaginary parts demonstrate a higher

percentage of parenchymal contribution to the respiratory elastance. Negligible chest wall

contribution to the inertive properties of the respiratory system is evident in the high-

frequency imaginary parts of the impedance spectra. Apart from the data points corrupted

by cardiac artifacts the model fitted all impedance data well with only a slight systematic

fitting error at the low-frequency real parts.

3.4.2 Repeated measurements in individual animals

The results of the repeated tests are summarized in Figure 3.2. Repeated measures of

respiratory, chest wall, and lung mechanics were highly reproducible with none of the

parameters showing statistically significant differences. Since repeated tests did not reveal

any significant change, w e used the results of the three tests to calculate the lung and chest

wall contributions to the total respiratory parameters. Figure 3.3 demonstrates these ratios

for the airway and tissue parameters. The frequency independent resistive and inertive

components of Zrs are fundamentally determined by Zl (90±0.8% and 100±1.6% for R and

I, respectively). The lung contributed a slightly higher part (58±1.5%) to the respiratory

system H, while it comprised a minority of the respiratory system G (33±1.0%).

73

o 3C?

a o o

750 -i

500

250 -

0

-500 -

-1000 -

-1500

Frequency (Hz)

Figure 3.1: Real (R) and imaginary (X) parts of a representative Zrs (•), Z w (•), and Z1(A) spectra (mean ±

SD) in intact chest condition. Hollow symbols represent data points corrupted by the cardiac noise and are

excluded from the model fit. The solid lines denote the corresponding model fits.

3.4.3 Comparison of the lung mechanics obtained in closed and open chest conditions

Figure 3.2 demonstrates that we found statistically significant differences in the airway

parameters when they were determined in situ and in the open chest. Open chest Raw was

13% lower (p<0.005) while law exhibited a 1 4 % increase (p<0.01). G showed good

74

agreement between the in situ and open chest condition with a slight and not statistically

significant increase of 9%. Chest opening caused a statistically significant 4 0 % decrease in

H (p<0.05). This decrease in H with a constant G resulted in a significant 3 4 % increase in

ri (pO.OOl).

Table 1. Sample size predictions from the values obtained in intact chest condition

Raw

G

H

MeaniSD

68±10 cmH20.s/L

401±72 cmH20/L

2665±363 cmH20/L

Expected SD of

change

12.5 cmH20.s/L

77.4 cmH20/L

578 cmH20.s/L

Sample no. required for change of

10%

35

38

47

20%

12

13

16

50%

5

5

6

Table 1. Sample size predictions to estimate 10%, 2 0 % , and 5 0 % change in Raw, G, and H. Mean and S D

values were obtained from the intact chest parameters. Expected S D of change was calculated from the repeated

Zl measurements. Mean R a w was corrected for the resistance of the E T tube and connecting tubing.

75

&

zuu -

100 -

9 9

• '""

. . B 9

i "*

. 1 1

l.O-i

d 0.5-

0.0

i I i i : i i i i i

1500 -i

^ 1000 o

o O 500

0.50 -i

p- 0.25

0.00 " T " 1st

— r — 2nd

~T~ 3rd

5000 -i

S &f 2500 -

1st 2nd 3rd

Figure 3.2: Respiratory (•), lung (•), and chest wall (A) parameters for the successive intact chest

measurements. Open chest parameters (mean±SEM dashed and dotted lines respectively) are also shown. *

p<0.05, #p<0.01, § p<0.005 and J pO.OOl.

76

Raw Rw law Iw Gl Gw HI Hw

Figure 3.3: Lung (hatched bars) and chest wall (white bars) contributions to the airway and tissue

parameters of Zrs.

3.4.4 Sample numbers for non-invasive, repeated and for single invasive studies

Based on the variability of the intact and open chest Zl data, w e calculated the required

number of rats to detect a 10%, 2 0 % , and 5 0 % change in Raw, G, and H. Tables 1. and 2.

show these sample size predictions for the intact and open chest, respectively. In all

parameters, the open chest condition required significantly greater (approximately 2-fold)

group numbers to show 1 0 % and 2 0 % change. This difference in the required number of

rats between groups is more pronounced (5-fold) if the primary variable of interest is G.

77

Table 2. Sample size predictions from the values obtained in open chest rats

Raw

G

H

Mean±SD

46±9.3 cmH20.s/L

439±146 cmH20/L

2899±465 cmH20/L

Sample no. required for change of

10%

83

212

50

20%

25

64

16

50%

7

17

5

Table 2. Sample size predictions to estimate 10%, 20%, and 5 0 % change in Raw, G, and H. Mean and S D

values were obtained from the open chest parameters. Mean Raw was corrected for the resistance of the ET tube

and connecting tubing.

3.4.5 M c h responses detected non-invasively

To examine the suitability of the present non-invasive method to detect constrictor

responses, aerosolized M c h was administered. Table 3 demonstrates these responses.

Parenchymal parameters exhibited statistically significant increases with percentage changes

from baseline of 393±109%, 111±27%, and 120±25% for G, H, and rj, respectively. N o

statistically significant changes in airway parameters occurred, however there was a

tendency in R a w to increase (56±27%, p<0.1), while law tended to decrease (-24±24%).

Chest wall parameters showed no statistically significant change following Mch-induced

constriction with percentage changes from baseline of-29±29%, 14±18%, -2.6±5.7%,

7.6±13.7%, and 3.1±10.4% for R, I, G, H, and T| of the chest wall, respectively.

78

Table 3. Effect of aerosolized Mch on in situ airway and parenchymal parameters.

Rat#

1

2

3

4

5

6

Mean

±SEM

Raw

cmH20.s/L

C

78

72

65

76

53

72

69.3

±3.7

MCh

250

333

206

44

90

61

164

±48

law

cmH20

C

0.57

0.68

0.73

0.68

0.78

0.63

0.68

±0.03

.s2/L

MCh

0.35

-0.19

0.54

0.85

0.66

0.87

0.51

±0.16

G

cmH20/L

C

437

261

439

369

435

451

398

±30

MCh

1800

2256

3297

1734

702

1344

1856

±358*

H

cmH20/L

C

2804

2367

2827

2264

2700

3035

2666

±120

MCh

5122

6967

7909

5153

3780

4330

5544

±646§

*1

C

0.16

0.11

0.16

0.16

0.16

0.15

0.15

±0.01

MCh

0.35

0.32

0.42

0.34

0.19

0.31

0.32

±0.03#

Table 3. Peak responses to inhaled M c h in individual rats. Statistically significant differences from baseline: *

p<0.005, § pO.0001, # p<0.0005.

3.5 Discussion

In the present study the low-frequency forced oscillation technique has been modified

[5,26,28,135,136] to characterize airway and parenchymal mechanics in rats non-invasively.

As the current method allows long-term investigations, in which individual rats can be used

as their own controls to study the development (or resolution) of a disease process, w e

studied the individual and group variability of lung mechanics over a two-week period. The

results of this study demonstrated that our technique requires significantly smaller group

sizes to detect small, but clinically significant (10-50%) changes in airway or parenchymal

mechanics when using the rats as their own controls than to detect the same change from

independent groups. To assess the potential of our method to observe changes in the airway

and parenchymal mechanics, we determined the lung response to inhaled Mch. In agreement

with previous findings we found a significant parenchymal contribution to the overall lung

constriction [5,138,139]. W e should note, however, that enhanced inhomogeneity of the

peripheral airways during constriction may add a significant virtual (not tissue origin)

component to the pulmonary G, and hence r\ [5,28]. Previous work by Petak et al. [5] has

demonstrated that even in the presence of moderate, methacholine induced constriction that

79

the changes seen in H are unaffected by peripheral airway inhomogeneities. Thus, the

marked and statistically significant increase in H may indicate a real alteration in the

mechanical properties of the parenchymal constriction. In contrast, the substantially greater

lung responses in G and r\ is likely to be influenced by peripheral airway inhomogeneities

r, [5,28].

Rats are commonly used to characterize the development of allergic respiratory diseases,

such as asthma, since the response of the lung to exogenous constrictor agents and allergic

inflammation in this animal model is well documented. Furthermore, recent studies on

different animal species have demonstrated the importance of the tissue resistance in

detennining baseline lung mechanics and the lung responses to various stimuli

[5,26,28,135,136,138-143]. Most of the previous techniques to measure parenchymal

resistance (e.g. alveolar capsule) in rats are invasive [138,140-142]. Although these

techniques provide details of the parenchymal mechanics, they are not suitable for repeated

measurements on the same animal. Therefore, either large number of animals were needed

or the statistics were not convincing to detect slight long-term changes in lung mechanics

that were present. Indeed, w e statistically calculated in the present study that the baseline

variability in parenchymal mechanics leads to a study using impracticably high numbers of

rats (N=64) to detect 2 0 % change in G. Recently, Oostveen et al developed a technique

to collect Zrs data in conscious rats non-invasively, using a two-compartment whole body

box [144]: Using their method, Hessel et al. examined the long-term reproducibility of Zrs

data in unanaesthetized mice [145]. The frequency range used in this technique (16-208

Hz), however, is unable to provide information about the resistive properties of the

respiratory tissues.

3.5.1 Validity of Pes measurement

A critical factor in measurement of lung mechanics in the intact chest is the ability of the

esophageal catheter system to accurately represent pleural pressure (Ppl). Generally, the

occlusion test is used to assess the validity of this hypothesis, in which the Pes/Ptr ratio is

recorded during spontaneous respiratory efforts or as a result of an externally applied force

on the chest [130] while the trachea is occluded. D u e to the absence of airflow during these

maneuvers, the Ptr is expected to be equal the alveolar pressure (Palv). Since changes in Pes

80

are also reflected in Palv, the slope of Pes/Ptr is expected to be one. In this study the

esophageal catheter was positioned based on the criteria of a smooth Pes trace with minimal

cardiac artifact and a good agreement between Ptr and Pes during airway occlusion. Indeed,

w e found a Pes/Ptr ratio of 0.97±0.02 indicating an accurate estimation of the average Ppl

in the rats during control condition.

While it is generally accepted that Pes can be used to estimate Ppl in control state, the

relevance of this conclusion during induced constriction is not evident. Suki et al. measured

MCh-induced responses in intact and open chest dogs [143]. while they found a

significantly greater lung response in the open chest, they also concluded following a

detailed analysis of the Z w spectra in control and constricted states that Pes measurement

can be used to partition Zrs to Zl and Z w even during induced constriction. In the present

study, analysis of the Z w parameters revealed good agreement in control condition and

following M C h challenge, confirming that Zrs partitioning is likely to be correct via Pes

measurements in rats even during MCh-induced constriction.

3.5.2 Lung parameters

The R a w values obtained in the present study from the intact chest Zl data were in all cases

significantly higher than those reported previously in tracheotomized rats

[5,28,135,138,140,141]. The fact that a significant part of our R a w values are attributed

to the flow resistance of the E T tube (98 of 166 cmH20.s/L) explains the majority of the

discrepancy between our present results and previously reported R a w values. Indeed if the

contribution of the E T tube to R a w is taken into account, our results are in good agreement

with those obtained using a similar measurement technique [5,28,135]. The systematically

lower R a w values derived by Nagase et al [138,140,141] with alveolar capsules during

tidal breathing (ranging from 24 to 35 cmH20.s/L) can be explained by the higher average

lung volume present during their measurements.

Lutchen etal. [28] and Petak et al. [5] determined G and H by modelling the low-frequency

Zlo spectra in open chest rats, and found 30-50% lower values than those obtained in the

present study. This can be attributed to body weight and strain differences. Our Rti values

calculated from the intact chest Zl data at the ventilation frequency (ranging from 32 to 70

81

cmH20.s/L) are comparable to those reported in the literature previously by Nagase et al

(38-43 cmH20.s/L) [138,140,141] and Navajas et al. (35 cmH20.s/L)[142]. Calculating El

from the intact chest XI data at 1.5 H z revealed significantly higher values (2400 to 4057

cmH 20/L) than those in previous studies performed in open chest rats (1233 to 1530

cmH 20/L) [28,138,140,141,146]. Beside body weight and strain differences, this

discrepancy can be attributed to methodological differences as in the latter studies El was

determined in thoracotomized rats during tidal ventilation.

3.5.3 Lung and chest wall contributions

Hantos et al [135] measured low-frequency Zrs and Zl spectra in intact and open chest

rats, and found a minor chest wall contribution to the Newtonian resistive properties of the

respiratory system, whereas the chest wall conferred the majority of the frequency

dependence of Rl (analogous to G ) . These findings are in qualitative agreement with the

results of the present study. Previous studies examining lung and chest wall components of

the total respiratory elastance in rats reported a relatively low (32-35%) chest wall

contribution [135,147]. The results of the present study with H w contributing 4 2 % to Hrs

is consistent with these previous findings.

3.5.4 Comparison of in situ and open chest conditions

Although w e kept the lung configuration in the open chest as close as possible to that of the

intact chest, significant differences were found in the pulmonary mechanics. While no

change was observed in G, significant differences were seen in Raw, law, H, and r| between

the intact and open chest conditions. Suki etal. [136] investigated the differences in airway

and lung tissue mechanical properties in dogs when they measured in situ and open chest

conditions, and found small but significantly different airway and parenchymal properties.

In their study no change was found in G, while R a w and TJ tended to increase following

chest opening. They also found a statistically significantly higher law and lower H in the

open chest condition. The present study demonstrated a similar pattern with the only

exception of Raw, which decreased. The matching of lung mechanics between the intact and

open chest conditions is critically dependent on matching the respective lung volumes. If

the Ptp was overestimated resulting in an increased lung volume in the open chest condition

a decreased R a w may be expected. The fact that our P E E P levels were relatively low (1-2

82

c m H 2 0 ) makes this argument unlikely, however species difference may play a role:

relatively small absolute errors in the P E E P would lead to a larger relative error in the rat

than in dogs, because the Ptp in the latter is higher. Alternatively the differences seen could

be due to alterations in the ventilation distribution. With the chest-wall intact, the increase

in lung volume will occur in a more-or-less even fashion. However, without the resttaining

influence of the chest wall, most of the increase in volume may occur anteriorly, with little

increase in the volume of the posterior lung units. The resulting distortion of the lungs could

have variable and unpredictable effects on lung mechanics.

3.5.5 Conclusion

In conclusion, w e have adopted a method to characterize airway and parenchymal

mechanics in rats non-invasively. W e demonstrated that this measurement technique is

suitable for studying the progression of disease or long-term treatment effects with

relatively small group size. The chest wall has been shown to contribute significantly to the

overall tissue mechanics of the respiratory system, while it has a negligible effect on airway

mechanics. Therefore, the present technique can be used to estimate long-term changes in

airway mechanics without Pes measurement, whereas estimation of the transpulmonary

pressure (either in an open chest animal or Pes) is advocated to study long-term responses

in Rti.

83

CHAPTER 4: CHEST WALL AND NASAL

CONTRIBUTIONS TO LOW-FREQUENCY

RESPIRATORY IMPEDANCE IN INFANTS

4.1 Summary

In 2 separate populations the low-frequency respiratory system impedance (Zrs) was

partitioned into lung (Zl) and chest wall (Zw) compartments (Study 1: n=5, mechanically

ventilated patients undergoing cardiac surgery) or nasal (Zn) and lower respiratory system

(Zlrs) (Study 2: n=l 1, spontaneously breathing infants) compartments using an esophageal

or nasal catheter, respectively. In both populations an oscillatory signal (Study 1: 0.5 - 12

Hz, and Study 2: 0.5 - 21 Hz) was applied during a pause in breathing to obtain the

respective impedance spectra. In both studies the Zrs spectra could be reliably partitioning

into either Zl and Zw, or Zn and Zlrs components. A model of the lung containing an airway

and tissue compartment was then fitted to the spectra. The airway compartment consisted

of a frequency-independent resistance (R) and inertance (I), while the tissue compartment

was described by a constant-phase tissue damping (G) and elastance (H). The nose was

found to contribute approximately half of the airway resistance (44.6±4.9%; m e a n ± S E M )

and a majority of the respiratory system inertance (71.7±3.5%), while exhibiting negligible

influence on the frequency dependent tissue mechanics. The chest wall made no appreciable

contribution to the frequency independent properties of the respiratory system, but made

a significant contribution to the Grs and Hrs (38.5 ± 7.3% and 34.4 ± 7.4%, respectively).

In conclusion the infant low-frequency respiratory system impedance can be represented as

a series element containing Zn, Zl and Zrs compartments. Future studies investigating

changes in tissue or airway mechanics should use esophageal or nasal catheters, respectively

to increase the sensitivity of the tests.

84

4.2 Introduction

Current methodologies for ascertaining lung function in infants involve measuring flows and

pressure via the nose in sedated patients [7,38,47]. These tests include the nose and extra-

thoracic airways, the chest wall as well as the pulmonary airways and tissues and it is the

pulmonary component that is of primary interest.

The frequency dependence of the mechanical properties of the respiratory tissues is well

established [7,18,148]. The low-frequency forced oscillation technique has been shown to

simultaneously provide information on airway mechanics as well as characterize the elastic

and resistive tissue properties in animal [4,149] and human studies [7,96,116]. Numerous

studies have reported estimates of resistance and compliance of the respiratory system,

chest wall and the lungs during mechanical ventilation. However, despite the fact, that the

short apnoeic periods required for low-frequency impedance measurements can be easily

achieved in anaesthetized, paralyzed patients, only a few Zrs [150] data below 2 H z have

been reported in adults and no such data is available in infants or young children. Petak et

al [12] have demonstrated that non-invasive longitudinal measures of Zrs, Zl and Z w can

be obtained through the use of an esophageal pressure catheter and the low-frequency

forced oscillation technique in anaesthetized, paralyzed Brown Norway rats.

In infants undergoing lung function tests the nasal resistance (Rn) may account for a

significant proportion of the overall respiratory system resistance [120,151,152]. As infants

are preferential nose breathers in the first few months of life it is important to measure the

active respiratory circuit. Nasal resistance (Rn) [153,154] and impedance (Zn) [155] have

been well characterized in older children and adults, while relatively few studies have been

conducted in infants [120,151,152,156]. These studies have used a number of adaptations

of the existing rhinometric [151], esophageal [152,156] or forced oscillation techniques

(FOT)[120] to characterize the nasal resistance or impedance.

The aims of the present study were: 1) to adapt the technique developed by Petak et al. [12]

and to determine the airway and tissue properties from low-frequency Zrs, Zl and Z w

spectra in anaesthetized, paralyzed infants and young children; 2) to characterize the

frequency dependence, resistive and inertive properties of the nasal impedance; and 3) to

85

determine the relative contributions of the nose, chest wall and lung to low-frequency

respiratory system mechanics.

4.3 Methods

The study was conducted in 16 subjects separated into 2 populations. The partitioning of

the respiratory system into pulmonary and chest wall components was carried out in

patients undergoing cardiac surgery (n=5). The second part of the study to investigate the

nasal impedance characteristics was carried out in sedated infants (n=l 1).

4.3.1 Subjects

4.3.1.1 Determination ofZl and Zw

Five children undergoing cardiac surgery were studied. Anthropometric data are presented

in Table 4.1. None of the children had acute respiratory symptoms at the time of the

measurement and none had history of chronic respiratory disease. T w o children were

operated on for atrial septal defect (ASD), one for ventricular septal defect (VSD) and two

for transposition of the great arteries (TGA-VSD). The protocol was approved by the

institutional H u m a n Ethics Committee and written parental consent obtained.

Following oral midazolam premedication (0.75 mg/kg), the children were anaesthetized by

intravenous (iv) administration of fentanyl (5-15 ug/kg) and midazolam (0.1-0.2 mg/kg).

Muscle paralysis was accomplished by an iv bolus of pipecuronium bromide (0.1 mg/kg).

Maintenance doses of the anaesthetic and muscle relaxants were administered as needed.

The patient was intubated with an uncuffed endotracheal tube (3.5-6 m m ID) and the

children were mechanically ventilated with a Siemens Servo Ventilator 900C in constant

flow mode by setting the inspiratory/expiratory ratio (I:E) to 1:3. The ventilator frequency

was set to 18-28/min and a tidal volume (VT) of 8-12 mL/kg was applied. The children

were in the supine position throughout the study.

Respiratory pressure, end-tidal C02, the fraction of inspired oxygen (Fi02), and the dynamic

respiratory compliance (Crs) were monitored by a Siemens Servo Computer Module (Type

990) attached to the respirator. Arterial blood gases were analyzed (Radiometer A B L ™

505) and the percutaneous 0 2 saturation was continuously followed (Datex Satlite).

Electrocardiogram, arterial, central venous, pulmonary arterial and left atrial pressures were

86

displayed on a monitor of a data acquisition system (Hewlett Packard M996).

4.3.1.2 Characterization ofZn

This study involved 11 infants ranging from 8 months to 2 years recruited from the general

population. Respiratory history was determined from a parental questionnaire.

Anthropometric data on the infants are outlined in Table 4.2. At the time of testing all

infants had been free of respiratory infections for a period of at least four weeks. The infants

were sedated with an oral dose of choral hydrate (70-100 mg/kg) and were laid in the

supine position with the head supported in the midline with the neck slightly extended. The

experimental protocol was approved by the H u m a n Ethics Committee of the Princess

Margaret Hospital for Children. Parents gave written informed consent and were generally

present during the study.

Table 4.1. Physical and clinical data on the cardiac patients

Patient

1

2

3

4

5

Age (mths)

72

48

84

90

36

Weight (kg)

16

16

30.5

17.7

15.2

Height (cm)

114

106

133

124

95

Sex

M

M

M

F

F

Diagnoses

TGA-VSD

VSD

ASD

TGA-VSD

ASD

4.3.2Measurement apparatus


See Section 2.2.2.2 for full details.

The measurement set-up used to collect Zrs data on spontaneously breathing infants [7] was

modified to measure impedance in the intubated children. Preceding oscillatory

measurements the lungs were inflated to a pressure of about 30 c m H 2 0 to standardize the

volume history. During the apnoeic period, smaU-amplitude pseudo-random pressure

excitations were introduced into the trachea. The forcing signal contained thirty integer-

multiple components of the fundamental frequency 0.4 Hz, resulting in a frequency range

of 0.4-12 Hz. Four to six recordings were collected in intact-chest condition. Periods of

at least 2 min were interposed between the successive measurements. The impedance curves

were then ensemble-averaged.

87

Table 4.2: Anthropometric data

Patient

6

7

8

9

10

11

12

13

14

15

16

Age (mths)

9.3

14.5

10.9

12.5

18.1

8.4

19.8

18.1

18.3

21.0

25.2

Weight (Kg)

10

10

8.7

9.2

10.5

11.2

10.5

11

12.8

14.9

11

Length (cm)

67

70

71

71.5

75

76

79

80

80.5

87

90

History

Healthy

Healthy

Healthy

Healthy

Wheezy

Healthy

Wheezy

Healthy

Healthy

Healthy

Cough

To separate Zrs into Z L and chest wall (Zw) components in the intact-chest, the esophageal

pressure (Pes) with reference to the atmosphere was measured. The pressure drop across

the chest wall was monitored during mechanical ventilation, and the catheter was positioned

to obtain a smooth respiratory curve with ininimal cardiac noise. The mechanical impedance

of the respiratory system was calculated from the Ptr and V signals (Zrs = Ptr/V'). The

chest wall impedance was computed directly from the Pes and V signals (Zw = Pes/V1),

while Zl spectra were obtained by subtraction (Zl = Zrs-Zw).


See Section 2.2.2.1 for full details

Respiratory system input impedance (Zrs) were measured using the low-frequency forced

oscillation technique during a pause in breathing induced by the Hering-Breuer reflex as

described by Sly et al. [7]. Nasal pressure (Pn) was determined using a solid state miniature

tipped catheter ( M T C 5F; Drager, Best Netherlands) placed 4-5 cm into the nasal cavity,

slightly above the nasopharynx. The infant, loudspeaker and a pressure reference chamber

was inflated three times to a transrespiratory pressure of 20 cmH20 with passive deflations

between each inflation. Following the third inflation, the airway was occluded inducing the

88

Hering-Breuer reflex. In the apnoeic period, a low-frequency pseudo-random signal,

between 0.5 and 21 H z was driven into the infant's respiratory system by the loudspeaker.

Five to ten measurements of the respiratory system (Zrs) and the lower respiratory system

(Zlrs) were collected in each infant.

The corrected individual respiratory impedance spectra were then fitted to a linear model,

in the 0.5-15 H z range [26] including: R and I equating to the frequency-independent

resistance and inertance of the airways, respectively and G and H representing the

constant-phase tissue damping and elastance, respectively. The impedance of the nose (Zn)

was calculated from the respiratory parameters of Zrs and Zlrs (e.g. Rn=Rrs-Rlrs). Tissue

hysteresivity (r|) was calculated as G/H [137].

4.3.3Study protocol and analysis


Baseline oscillatory parameters were averaged for each patient (mean±SD) and the

individual contributions of Zl and Z w mechanics to Zrs mechanics calculated.


Lung function for each infant was expressed as the mean ± S D of the individual respiratory

parameters. In a sub-group of patients (n=4) the effect of introducing the nasal pressure

catheter was assessed. Linear regressions were used to quantify the frequency dependence

(Fd) of the nasal resistance and reactance. Group means and standard errors of the

mean(SEM) are reported for R n and Xn. Pearson Product Moment correlation was used

to examine the relative contribution of the nasal impedance to Zrs and its association with

growth. Significance was accepted at p<0.05.

4.4 Results

4.4.1 Determination of Zl and Z w

Representative Zrs, Zl and Z w impedance data (child #4 ) and the corresponding model fits

are demonstrated in Figure 4.1. The plateau evident at higher frequencies represents the

resistance of the airways and the chest wall. The domination of Zrs by Zl at these higher

frequencies indicates the major contribution of R a w to Rrs, with the chest wall playing a

89

negligible role. The negative frequency dependence of Rrs and Xrs at lower frequencies

represents the frequency dependant mechanical properties of the respiratory tissues and

chest wall. Negligible chest wall contribution to the inertive properties of the respiratory

system is evident in the high-frequency imaginary parts of the impedance spectra. Data

points corrupted by cardiac artifacts were ignored. The model fitted all impedance data well

with only a slight systematic fitting error at the low-frequency real parts. Individual

pulmonary and chest wall components of the total respiratory mechanical parameters are

shown in Figure 4.2.

20 -,

w

E o 5 CO

S. E o CD

15

10

5

0

50

40

30

20

10

0

0.6 n

0.4

0.2 -

0.0

Zrs

Zrs

ZL

•

v

o S

ZL

"fl" ~

Zw

Zw

a, E o CO

40

20

-20

250

200 -

g.,150 E 3. 100 X

50 -(

0

9

Zrs

Zrs

ZL

I o

ZL

Zw

Zw

Zrs ZL Zw

Figure 4.1: Representative Zrs (•), Zl (•) and Zw (•) spectra from child #4. Baseline spectra were averaged

(mean±SD). The corresponding model fits are shown as solid lines. Frequencies affected by cardiac noise or

it's harmonics have been excluded from the fit and are not shown.

While w e found a great scatter in the individual parameter values, which mainly reflect the

90

differences in body size, the relative contributions of the lungs and the chest wall to the

mechanical properties of the total respiratory system were uniform. The lungs determined

primarily the R and I parameters of the total respiratory system (90.8±4.3% and

109.3±16.4% for R and L respectively) while contributing a majority of the total respiratory

G (61.5±7.3%) and H (65.6*7.4%). The -q was highest for the chest wall (0.30±0.06), while

it was the lowest for the pulmonary system (0.20±0.03).

4.4.1 Characterization of Zn

Figure 4.3 illustrates an averaged impedance measurement and its corresponding model fit

for Zrs and Zlrs. Zn is also shown. The negative frequency dependence of Zrs and Zlrs

within both the resistive (Rrs) and reactive (Xrs) components represents the frequency

dependence of the respiratory tissues. The resistive plateau observed at higher frequencies

is associated with the frequency independent properties of the airways and nasal pathways.

The zero crossing point in Xrs represents the resonant frequency of the respiratory system

and is predominantly influenced by the inertance of the nasal cavity. The frequency

dependence (Fd) of the nasal resistance (Rn) and reactance (Xn) was assessed by

deterrnining the slope of the regression line across the frequency range (0.5-15 Hz) for each

child. The resistive properties of Zn are predominately frequency independent, with a Fd

of-0.03±0.03 (mean±SEM; range: 0.14-0.22). The reactive component exhibits a strong

linear frequency dependence across the whole spectra and predominantly represents the

inertive properties of the nasal cavity, with a Fd of 0.91±0.17 (range: 0.44-2.46).

91

40

b

20

*

~i »i»* **» * • « »

Frequency (Hz) 10

Figure 4.2: Individual respiratory, lung and chest wall mechanics (shown as symbols). Group mean data are also

shown (solid lines). The chest wall contributed a negligible amount to the Rrs and Irs and significantly influences

the tissue visco-elastic properties of the respiratory system.

Nasal impedance was found to contribute significantly to the airway properties of the

respiratory system, but to have a negligible effect on the tissue mechanics (Figure 4.4). The

lung and nasal pathways had approximately equal contributions to the airway resistance of

Zrs (55.4±4.9% and 44.6±4.9%; mean±SEM, respectively), while the nasal cavity

dominated the overall inertance of Zrs (71.7±3.5%). The lower respiratory system

contributed exclusively to the tissue mechanics of the respiratory system (93.4±2.1% and

95.3±2.1% for G and H respectively). The relative contribution of the airway and

parenchymal mechanics of both the lower respiratory system and nasal cavity to the overall

respiratory system mechanics was found to have no systematic variation with body length.

In a subgroup of 4 infants the effect of the nasal catheter placement on Zrs was assessed.

N o significant effect was found in any parameter (Table 4.3)

92

60 -,

40

•a 20-

20

W 1 -20

s o

-40 -

-60

-80 J

Frequency (Hz) 10

Figure 4.3: Representative spectra from an individual infant. Average Zrs (•), Zlrs (A) and Zn (•) spectra

(mean±SD) are shown and their corresponding model fits (solid line). Open symbols represent data points

corrupted by cardiac noise and are excluded from the model fit.

93

Table 4.3: Effect of nasal catheter placement of airway and respiratory tissue mechanics

Patient

2

4

6

10

Mean

(SEM)

Raw (cmH20.s/L)

Pre

18.29

(3.27)

8.99

(1.09)

14.04

(0.61)

19.91

(1.63)

15.16

(2.36)

Post

16.14

(0.54)

11.27

(1.20)

17.03

(0.50)

22.59

(1.71)

17.16

(2.66)

law (cmH20.s2/L)

Pre

0.12

(0.01)

0.07

(0.03)

0.11

(0.01)

0.16

(0.02)

0.11

(0.02)

Post

0.11

(0.003)

0.07

(0.01)

0.14

(0.01)

0.18

(0.02)

0.13

(0.02).

Grs (cmH20/L)

Pre

25.6

(2.18)

23.6

(5.82)

26.7

(2.77)

58.4

(8.19)

34.5

(9.2)

Post

31.3

(0.27)

22.3

(10.09)

26.1

(7.29)

48.8

(13.06)

32.4

(6.1)

Hrs (cmH20/L)

Pre

118.9

(1.8)

105.1

(5.3)

145.9

(4.0)

158.2

(8.60)

132.7

(12.7)

Post

122.3

(6.0)

102.1

(9.1)

134.6

(3.1)

140.4

(17.8)

128.8

(11.3)

Table 4.3: Individual and group infant data for pre- and post nasal catheter placement (Individual data;

mean(SD) and group data; mean(SEM)). N o significant difference was seen following catheter placement.

4.5 Discussion

The purpose of the present study was to: firstly, determine the low-frequency impedance

spectra of the nose, chest wall and pulmonary components of the respiratory system and

secondly, to calculate the contribution of these components to the constant-phase model

parameters of the Zrs. Respiratory system, chest wall and lung impedance data were

determined in five children, prior to the chest being opened for cardiac surgery, at

frequencies encompassing the spontaneous breathing rate during short apnoeic periods

introduced into the mechanical ventilation. Zrs and Zlrs spectra were determined in 11

spontaneously breathing infants, during a pause in breathing, induced by the Hering-Breuer

reflex, at a transrespiratory pressure of 20 cmH20. Furthermore these impedance spectra

were fitted to the constant-phase model [4] allowing simultaneous assessment of the airway

and tissue mechanics of the respiratory system and lung, the mechanical properties of the

chest wall, and the resistive and inertive properties of the nose. The overall contributions

of the nose, chest wall and the lung to the respiratory system were also quantified.

94

to

d.

f o

20

10-

0.2

0

50

25-

f 0.1

»:§:§sSj

f

0.0

150

100 -

50-

-

I T

J888HHfa«.

Zrs Zlrs Zn Zrs Zlrs Zn

Figure 4.4: Partitioning of respiratory system mechanics (open bar) into lower respiratory (diagonal bar) and

nasal Oiatched bar) mechanics. Group mean data (±SEM) are shown. The lower respiratory system contributed

approximately half of the resistance, but had a minor influence on the inertance of the respiratory system.

Respiratory system tissue damping and elastance were exclusively located in the lower respiratory system with

the nasal pathways having a negligible contribution to the tissue mechanics.

4.5.1 Partitioning of Zrs into Zl and Z w

4.5.1.1 Total respiratory impedances

It was demonstrated that, in agreement with spontaneously breathing infants [7,76] and

various animal species [4,5,149], the marked frequency dependence of the Zrs can be

described by a model containing a frequency independent airway and a constant-phase

tissue compartment. The constant-phase model of the lung has been shown to reliably

describe the mechanical properties of the respiratory airways and tissues [5], Lutchen et al

[27] demonstrated that in the presence of peripheral airway inhomogeneities, resulting from

severe constriction or inhomogeneous ventilatory distribution patterns, that a non-tissue

related component may be present in the model estimation of tissue damping (G).

95

Peripheral ventilation inhomogeneities may be present due to partial airway closures in our

supine sleeping children. To minimize this phenomenon in the present measurements of Zrs

a deep inflation before successive Zrs measurements was performed to ensure recruitment

of possibly atelectatic lung regions. Furthermore, the presence of a virtual G component

would be reflected in an unusually high respiratory system ri values, however this was not

the case in the present study (n,rs: 0.17-0.27). Therefore while the possibility of peripheral

inhomogeneities cannot be excluded it is likely that any component of G that is non-tissue

related would be negligible, thus it can be concluded that the airway parameters derived

from the low-frequency Zrs spectra represent the flow resistance of the bronchial tree while

the model parameters of the tissue compartment characterize reliably the viscoelastic

properties of the respiratory tissues.

The Crs values calculated in the present study at 0.4 Hz (Crs = {Hrs co/©"}"1: 8.42±3.63

mL/cmH 2 0 ) were almost identical with values derived from the height-matched values of

Hantos et al (8.34±3.74) [157]. In contrast the mean value of Ers in the present study

(141.1±68.6 cmH 20/L) were approximately three times those found by Lanteri et al.

[158](51.1±17.9 cmH 20/L) using the multiple linear regression ( M L R ) technique during

mechanical ventilation in cardiac patients of similar heights. In the present study lung

mechanics were determined at F R C , while in the study by Lanteri et al. [158] lung

mechanics were estimated from 20 second epochs of V and Pao during mechanical

ventilation. Since Ers has been shown to decrease with increasing lung volume the higher

values of Ers in the present study are expected [107,148,159].

Despite differences in methodologies the derived Rrs values from the present study compare

well to those reported in the literature. The Rrs data in the present study (Rrs=Raw+Rt:

20.3±10.5 cmH20.s/L ; Rti=G/coa) compares well to that of cardiac patients of similar

height (25.2±9.6 cmH20.s/L) [158] and to those derived from regression equations

determined from normal children (27.4±6.9 cmH20/L/s) [61].

4.5.1.2 Lung impedances

Numerous studies have examined lung mechanics in adults both before and after chest

opening [148,160]. Lanteri et al. [158] determined the lung mechanics of 9 children

undergoing cardiopulmonary bypass surgery, 3 of which fall within the age range of the

96

present study. The mean height matched El of 24.4±10.9 cmH 20/L was lower than the

mean El in the present study (86.5±31.5 cmH20/L) and can be attributed to the lower mean

lung volumes used in the present study, with lung elastance reported to decrease

proportionally with increasing lung volume. [2].

4.5.1.3 Contribution qfZlandZw to Zrs

Reported contributions of the chest wall to respiratory system mechanics have been varied.

Barnas et al. [108] reported approximately equal contributions of the lung and chest wall

to the respiratory system elastance, independent of frequency, with the chest wall

contribution increasing slightly with decreasing tidal volume. The investigators also

reported chest wall contributions in excess of the pulmonary resistance at low frequencies.

Studies in adult humans have shown the frequency dependent characteristics of Zrs to be

predominantly due to the chest wall particularly at low frequencies [96]. Approximately

equal contributions of pulmonary and chest wall elastance obtained by small amplitude

oscillations at F R C were reported by Hantos et al [96]. Barnas et al. [148,160] reported

slightly lower chest wall contributions of approximately 40- 4 5 % and 17-23% of respiratory

system elastance and resistance, respectively at a Vt of 250mL. The differences between

these studies most likely relate to the lung volumes at which measurements were made

[148,160]. In the present study the chest wall contribution to the respiratory system tissue

damping and elastance was 38.5±7.3% and 34.4±7.4% respectively. Papastamelos et al

[65] demonstrated increasing chest wall contribution up to 3.5 years of age. In infants

underl year of age the chest wall comprised approximately 2 5 % of the total respiratory

system elastance, this increased to 4 5 % in infants older than 1 year of age. Sharp et al

[161] reported chest wall elastance approximating 4 0 % of Ers. This contribution was found

to be constant in children between 3 and 18 years of age. In a group of five infants Nicolai

et a/. [18] reported that the chest wall elastance was frequency dependent, but that the

contribution of the chest wall to the respiratory system elastance did not change with

frequency. The contribution reported by the investigators (19±6%) is lower than that in the

present study and is most likely related to the older patients in the present study and the

increased E w resulting from small amplitude oscillations. Lanteri et al. [158] determined

lung mechanics in 9 children undergoing cardiopulmonary bypass surgery with the chest

wall contributing approximately 3 0 % of Ers, of these 9, 3 are within the age range of the

97

present study with the chest wall contributing 37.3±19.9% to Ers and these values are

comparable with those of the present study.

4.5.2 Partitioning of Zrs into Zn and Zlrs

The present study has demonstrated that the low-frequency forced oscillation technique can

be adapted to obtain reliable estimates of nasal impedance, that the C P M can be applied to

Zrs and Zlrs and the resultant mechanics of the infant nose determined. Measurements of

nasal resistance using rhinometry are well established in adults and older children [153,162-

164]. The forced oscillation technique has also been applied in the adult patient group using

measurements at the mouth and nares [165], utilizing the Valsalva manoeuvre [166] or by

measuring impedance at the mouth with the nose open and closed [155]. These techniques

have proved difficult to translate into an infant population. Stocks and Godfrey have

described an adaptation of the posterior rhinometry technique for use in infants [151], while

Desager et al. [120] used an adaptation of the F O T to quantify the frequency dependence

of the nasal impedance above 8 Hz. The present study is the first to characterize the low-

frequency characteristics of the nasal impedance and its influence on the total respiratory

input impedance in infants.

The acute affects of the nasal catheter were examined in a sub-group of patients. No

significant effect was seen in any respiratory parameter (Table 4.3). This is in contrast with

the results previously reported by Stocks et al. [167]. In a group of 4 Caucasian and 3

black preterm infants the author reported a mean increase of 1 0 1 % and 5 0 % respectively

when a nasogastric feeding tube (NGT) was passed through the larger nostril. The change

was more pronounced if the N G T was passed through the smaller nostril. In both studies

the catheter size was identical (FG5). The differences can be attributed to the size of the

infants at the time of testing. In the previous study [167] all of the infants were preterm and

were studied in the first 4 weeks of postnatal life, whereas the current study was carried

out in term infants ranging from 8 months to 2 years postnatal age. Hence the cross-section

area of the nose in the present would be considerably larger and the effects of the catheter

accordingly reduced.

4.5.2.1 Contribution ofRn to Rrs

In the present study R n was found to contribute 44.6±4.9% (mean±SEM) to Rrs. This

98

compares well to previously reported data of 4 9 % using posterior rhinometry [151] or 2-

4 9 % with oscillations above 8 H z [120] The influence of the nasal pathways to the total

respiratory damping (Gn/Grs) was negligible (6.6±2.1%). Desager et al [120] reported

minimal resistive frequency dependence in both normal and asthmatic children with a slight

negative dependence in children with a higher Rn. The authors however only examined

frequency dependence at 24 and 48 Hz. Using a lower frequency range of 3-15 H z Fullton

et al. [165] found the frequency dependance of the nostrils in parallel to be miiiimal in

normal adults. In the present study there was no significant resistive frequency dependance

regardless of symptomatic history although the numbers are small (healthy n=8,

symptomatic history n=3). The frequency dependance of the respiratory system resistance

predominantly results from the mechanical behaviour of the pulmonary tissues [7,26]. The

lack of frequency dependance, in R n indicates that the resistive properties of the nose can

be attributed solely to the nasal airway and that any frequency dependant resistive

component are negligible.

4.5.2.2 Contribution ofXn to Xrs

The nasal cavity provided the dominant contribution to the inertive properties of the

respiratory system (In/Irs = 71.7±3.5%; mean±SEM) while having a negligible influence on

the elastance of Zrs (Hn/Hrs = 4.7±2.1%). Previous studies have not attempted to

determine the influence of X n on Xrs. Desager et al [120] reported a positive frequency

dependance of X n with mean values of 0.9±3.7 cmH20.s/L at 24 H z and -0.1±1.9

cmH20.s/L at 48 Hz. While a direct comparison cannot be made, these results are

qualitatively similar to those reported in the present study.

In a cross-sectional study Stocks and Godfrey [151] found that the absolute values of Rn

decreased monotonically from birth to the end of the first year of postnatal life when plotted

against thoracic gas volume (TGV). However the percentage contribution of R n to R a w

remained relatively constant against T G V . This later result was reflected in the present

study. The percentage contribution of R n to R a w in the present population was relatively

constant with length in the population studied. Similarly the contribution of In to Irs was

also constant with length. This indicates that as in infant grows, the rates of change of the

respiratory system, lower respiratory system and the nose are approximately equal.

4.5.3 Conclusions

In conclusion, the results of the present study demonstrate that low-frequency mechanical

impedance spectra of the respiratory system, lung and chest wall can be reliably collected

in anaesthetized, paralyzed children prior to cardiac surgery. In addition the low-frequency

F O T can be used to partition the Zrs into nasal and lower respiratory components.

Partitioning of the lung and chest wall mechanics in the intact chest via Pes measurement

revealed the dominance of the lungs in the total respiratory mechanics in children. The

application the Zrs, Zl and Z w spectra to the constant-phase model allowed the

simultaneous assessment of the airway and lung parenchymal mechanics during surgery,

thus allowing any changes in the mechanical state of the lung to be monitored. The

quantification of the nasal impedance demonstrated the significant contribution of R n to

Rrs, with the nasal inertance dominating the respiratory system reactance. The ability to

partition the respiratory system into either lung and chest wall compartments or nasal and

lower respiratory system compartments will provide a greater degree of sensitivity to

diagnostic tests that wish to assess changes in respiratory system mechanics.

100

101

CHAPTER 5: DEVELOPMENT OF AIRWAY

AND RESPIRATORY TISSUE MECHANICS IN

HEALTHY INFANTS

5.1 Summary

Low-frequency respiratory impedance (Zrs) was measured by applying a forcing signal,

between 0.5 and 21 H z at a transrespiratory pressure of 20 c m H 2 0 , in a cross-sectional

study of 37 normal infants. A lumped parameter model containing airway (resistance (Raw)

and inertance (law)) and a respiratory tissue (tissue damping (G) and elastance (H))

compartments was fitted to the individual Zrs. Forced expiratory volume in 0.5 seconds

(FEV0.5) was determined using the raised volume rapid thoracic compression technique.

Multivariate regression analysis was used to analyze the relationships between the

mechanical properties of the airways and respiratory tissues and length. W h e n plotted as a

function of length, both airway and tissue parameters showed a decreasing quadratic

relationship with increasing length. FEV0.5 showed an increasing cubic relationship with

length. A family history of asthma was found to have a negative effect on Raw, H and

FEV0.5. The rate of change in airway mechanics was slower than the tissue mechanics

supporting the concept of dysanaptic growth within this population.

102

5.2 Introduction

The anatomical makeup of the respiratory system changes dramatically during the first years

of life. Developmental changes in the airways, chest wall and parenchyma occur, leading

to rapid transformations of the respiratory system mechanical properties. Previous studies

have shown consistent increases in spirometric indices until early adulthood, followed by

a slow decline in lung function with increasing age [168,169]. Standard spirometry is not

suitable for children younger than 6-7 years of age and hence cannot be used to describe

mechanical changes in the respiratory system of infants or young children. The rapid

thoracic compression technique (RTC) and more recently the raised volume rapid thoracic

compression ( R V R T C ) technique have been used to describe the changes in FEVi, F V C

and forced flows in sedated infants with respect to age, following administration of

broncho-active agents and in disease [38,47,131,170,171]. The interrupter technique has

been used to partition the respiratory system resistance (Rrs) into airway resistance (Raw)

and a visco-elastic tissue component and also to calculate quasi-static compliance [61]. The

R T C and R V R T C techniques, are limited to providing information on spirometric changes

in the age groups studied. Techniques that measure pressure and flow at the airway

opening, are limited to the inspiratory phase of respiration or in intubated patients under

general anaesthesia [61] and provide limited information.

A number of investigators have used the FOT to determine the mechanical properties of the

respiratory system in spontaneously breathing children using frequencies well above the rate

of spontaneous breathing (2-6 Hz) [102,157]. T o adequately describe the mechanical

properties of both the airways and the tissues, the oscillatory signal must incorporate

frequencies below that of the spontaneous breathing rate, the frequency dependence of the

resistive and elastic properties of the lung at low frequencies has been well established

[96,172]. Sly et al [7] recently described a technique that used the Hering-Breuer reflex

to induce a pause in breathing in sedated infants allowing reliable low-frequency respiratory

impedance spectra (Zrs) to be collected and partitioned into airway and tissue parameters.

The present study was performed to measure airway and lung tissue mechanics in healthy

infants in the first two years of life. Since both the R V R T C and low frequency forced

oscillation techniques were used in this study, w e were able to determine h o w oscillatory

and forced expiratory parameters reflected the changes in mechanics with growth.

103

5.3 Methods

5.3.1 Subjects

This was a cross-sectional study involving 36 infants (15 male, 22 female) of ages ranging

from 7 weeks to 2 years, recruited from the general population. Information regarding

parental smoking status (maternal, paternal and both parents: n=3,5 and 5, respectively) and

history of asthma in immediate family members (negative, positive and unknown: n=21,14

and 2 , respectively) were obtained. Twelve infants had their lung function measured on two

separate occasions. Forced expiratory data were obtained on 28 infants, 8 of w h o m had

repeated lung function measurements (total observations=36), while F O T data were

collected on 28 infants, with 6 having repeated lung function (total observations=34). In 23

infants both FEV0.5 and F O T were obtained. Infants were free of history of asthma,

bronchiolitis or bronchitis and had fewer than 3 occurrences of lower respiratory illnesses

per year of life. Anthropometric data on the infants are outlined in Table 5.1. The infants

were sedated with an oral dose of choral hydrate (70-100 mg/kg) and were laid in the

supine position with the head supported in the midline and the neck slightly extended. The

experimental protocol was approved by the Human Ethics Committee of the Princess




5.3.2.1 Forced oscillation technique

See Section 2.2.2.1 for full details

Respiratory input impedance spectra (Zrs) were measured by using the low-frequency

forced oscillation technique during an apnoeic pause induced by the Hering-Breuer reflex

as described by Sly et al. [7]. Three inflations were applied to a circuit containing the infant,

the loudspeaker and a pressure reference chamber. At the end of the third inflation, the

airway was occluded at a transrespiratory pressure (Ptr) of 20 c m H 2 0 to induce the Hering-

Breuer reflex. In the resulting short apnoeic period, a low-frequency pseudorandom signal,

containing 16 frequency components in the 0.5-20 H z range, was driven into the infant's

respiratory system by the loudspeaker. Measurements were 6 s in length and examined for

leak or respiratory efforts, with corrupted recordings being excluded. Three to ten

104

measurements of the uncorrected input impedance of the respiratory system were collected

in each infant. The impedance of the dead space was considered as a lumped shunt

impedance in parallel with the respiratory system, and the corrected respiratory impedance

(Zrs) was calculated as described previously [7]. A linear model of the respiratory system

consisting of an airway and tissue compartment [26] was then fitted to the Zrs data, in the

0.5-15 H z range. The airway compartment contained a frequency-independent resistance

(Raw) and inertance (law), while the tissue compartment included a constant-phase tissue

damping (G) and elastance (H). Tissue hysteresivity (T|) was calculated as G/H [137].

Scatters in the anthropometrical data and the respiratory mechanical parameters were

expressed in S D values.

5.3.2.2 Raised volume rapid thoracic compression technique


Forced expiratory data were obtained over an extended volume range using the technique

described by Hayden et al. [131]. The infants were inflated three times to a transrespiratory

pressure of 20 cmH 2 0 with passive deflations between each inflation. Utilizing a jacket

connected to a positive pressure reservoir, a compression force was applied to the thorax

and abdomen after the third inflation. The compression force was standardized to transmit

a pressure of 20 cmH 2 0 to the airway at end-inflation resulting in an airway opening

pressure of 40 cmH 2 0 [131]. Forced expiratory flow was determined, integrated and

volume-time curves were produced. FEV0.5, FEV0.75 and FEVi were calculated from the

forced expiratory volume-time curves. Three to ten technically correct measurements were

obtained from each infant.

105

Table 5.1: Anthropometric data

Patient #

1 2

2 Visit 2

3 4

4 Visit 2

5 6

6 Visit 2

7 7 Visit 2

8 8 Visit 2

9 10

10 Visit 2

11 12 13

13 Visit 2 14 15 17

17 Visit 2

18 19

19 Visit 2 20

20 Visit 2 21 22 23 24 25 26

26 Visit 2 27

| 28 ! 29

30 31 32 33 34 35 36

36 Visit 2

Gender

F M M F F F F M M F F F M M M F M M M F M M M M F F F F F F F F M F F F M F F F F M F F M M M

Age (weeks)

26.7

9.4 27.6

13.7

48.7

79.9 37.0

16.6

46.0

13.3 33.4

12.6 25.6 84.0 18.0 35.0

13.9 31.7 12.6

33.7 16.3 17.6 22.4

54.6 67.0 30.3 50.6 52.6 79.6 58.1 50.1 41.7 28.9 41.6 65.9 89.1 44.9 50.7

53.9 59.1 69.3 92.7

84.9 86.0

68.3

94.3 35.1

Height (cm)

70.0

58.0 68.0

58.0

70.0

81.0 67.0

60.0 71.0

61.0

74.0

55.0 62.0 87.0 64.0 67.0 64.5 65.0 65.5 76.0 66.0 66.0 67.5

80.5 80.9 68.5 72.8 77.0 86.0 70.0 71.5 72.0 73.0 73.0 74.0 79.0 76.0 76.5 77.0 78.0 78.9 81.0 81.7 82.4 82.5 83.0

70.0

Weight (kg)

8.9 6.3 9.6 6.1 7.9 10.4 10.0 6.4 8.7 6.0 10.0 4.2 6.7 14.9 6.9 8.3 7.5 7.4 7.6 11.3 8.9 7.9 8.3 11.8 12.5 8.2 9.8 10.8 13.2 10.0 9.2 9.6 10.0 8.4 10.2 11.5 10.5 10.9

8.2 12.0 11.4 11.2

11.2 11.3 14.4 12.5 10.8

RVRTC/FOT

y/y y/y y/y y/n n/y n/y y/y y/y n/y y/y y/y y/n y/y n/y yfy y/n y/y y/y n/y y/y y/n y/y y/n y/n y/y n/y y/y y/n y/n y/y y/y ym n/y n/y y/y y/n y/y y/n y/y y/n n/y y/y n/y y/y n/y y/y y/n

5.3.3 Study protocol and analysis

Baseline measurements using the low-frequency F O T and R V R T C techniques were

obtained from each infant. Multilevel models were used to examine the relationships

between lung function (natural logarithms of G, H, R a w and FEV0.5) and length (Lt). These

106

models accounted for the non-independence of the data due to the repeated measurements

for each child and repeated visits for some children. Initially, models with three levels

(individual observations, within separate visits and within children) were fitted, but variation

at the second level (that is, covariation between visits for the same child) was not

statistically significantly different from 0 after accounting for differences in Lt, so the model

was collapsed into a two level model. Lt (centred about the mean), Lt2 and Lt (after

centring) were potential covariates in all models. This allows a wide range of potential

functional relationships between Lt and response, varying from a straight line (coefficients

for Lt2 and Lt3 both 0) to a cubic curve with two turning points. Although a wide range of

other functional forms could have been considered, it is very unlikely that in a data set of

this size that a useful discrimination could have been made. Gender and family history of

asthma were considered as covariates to examine their influence. Smoking status was not

included due to small numbers. Family history of asthma was treated as a binary variable. To

examine the relationships between oscillatory parameters and FEV0.5 linear regression

models were used. As the oscillatory parameters are volume dependent and FEV0.5 is a

measure of volume, the oscillatory data were corrected for volume (using Lt as a substitute)

by multiplying oscillatory parameters by length. Four separate models were fitted, each

potentially including one oscillatory parameter and its square. Multiple regression models

including all oscillatory parameters and their squares were also fitted. A s the interactions

HZRaw and G/Raw were of interest, the inverse of R a w (length-corrected) was used rather

than Raw*Lt. The hysteresivity (rj) was also included as a potential predictor.

5.4 RESULTS

5.4.1 Low-frequency respiratory system mechanics

The relationship between natural log (Ln) of the oscillatory parameters and length was best

described by a quadratic equation (Table 5.2). M e a n individual values for Raw, G and H,

the fitted regression lines and 9 5 % confidence intervals are plotted in Figure 5.1. The mean

coefficients of variation within subjects were similar for all oscillatory respiratory

parameters (11.3±0.8%, 22.4±2.6%, 13.9±1.2% and 9.5±0.7%, for Raw, law, G & H

respectively: mean±SEM, Table 3). r| was found to have no significant systematic variation

with length (0.27±0.01) and was not modelled. As law is primarily influenced by the

extrathoracic airway it was not examined further. T o examine the concept of dysanaptic

107

growth in this population the rate of change in airway and tissue mechanics was examined

[61], by regressing Ln(Raw) against Ln(G) and Ln(H). Figure 5.2 demonstrates the

relationship between Ln(Raw) and Ln(G). The slopes of the regression equations were

comparable (0.47 and 0.49 for G and H, respectively).

5.4.2 RVRTC

The relationship between the Ln(FEV0.s) and length was best described by a cubic equation

(Table 5.2). Figure 5.3 shows mean individual FEV0.5 values against length, the fitted

regression line and 9 5 % confidence intervals. The mean intra subject coefficient of

variation (CV) was 7.6±0.7%. (Table 5.3).

Table 5.2: Regression Equations for the natural log respiratory mechanics

Equation

Ln(FEV0.5)=-27.718+1.329*Lt-0.0184*Lt2+(8.699*10-5)*Lt3

Ln(Raw)= 1.98 - 0.0513*Lt + (5.137*10-4)*Lt2

Ln(Iaw)= 0.0388 - 0.0293*Lt

Ln(G)= 9.723 - 0.1074*Lt + (iJS^W4)*^

Ln(H)= - 4.127 + 0.0859*Lt - (9.93*10"4)*Lt2

Variance

Within Child Between Children

0.01307

0.01718

0.14280

0.02854

0.01155

0.03866

0.05678

0.19830

0.05115

0.05399

108

^ 45-

30 -

15 -

80 90

Length (cm) Figure 5.1: Airway resistance (Raw), tissue damping (G) and tissue elastance (H) against length. Oscillatory

lung function parameters were fitted to a quadratic equation of length (solid line) , the long dashed line

represents the 9 5 % confidence intervals for the individual data. Mean measurements for each child are also

shown (•).

109

5.4.3 Comparison between F O T and R V R T C

All length corrected oscillatory parameters were well correlated with FEV0.5 (Regression

analysis R2: 26.8%, 16.5%, 62.9% and 72.8% for Raw, law, G and H, respectively).

FEV0.5 was best predicted by length-corrected elastance (H*Lt) and airway resistance

(Raw*Lt) and their squares ((H*Lt)2 and (Raw*Lt)2, respectively) (FEV0.5 = 865.4-

0.066(H*Lt)+1.9*10-6(H*Lt)2 -0.25(Raw*Lt) + 8.25*10"5(Raw*Lt)2; R2=79.3%,

p<0.0001). The addition of further parameters or interactions between parameters did not

significantly improve the model fit.

6 1

4 -

I 3

4

L n G (cmFfcO/L) Figure 5.2: Ln (Raw) vs. Ln (G). Solid dots (•) represent the mean Raw and G data in each child; solid line,

model best fitting these data (slope = 0.47).

5.4.4 Factors influencing lung function parameters

Gender was not a significant covariate within the present study. Family history of asthma

was a significant covariate for Raw, H and FEV0.5 (Raw: p<0.02, H: p<0.01 and

FEV0.5:p<0.005) and caused a worsening of lung function.

110

1 r

50 60 70 80 90

Length (cm) Figure 5.3: FEV0.5 vs. length. The relationship between FEV0.5 and Lt was best described by a cubic equation

(solid line), the long dashed line shows the 9 5 % confidence intervals for the individual data Mean measurements

for each child are also shown (•).

5.5 Discussion

The results of the present study show the mechanical properties of the respiratory system

over the first two years of life. The information obtained has been utilized to produce cross-

sectional reference values for the mechanical properties of the airways and the respiratory

tissues separately in sedated infants.

5.5.1 FOT

Previous data examining the changes with age in respiratory compliance (Crs) and R a w

have been determined using passive flow-volume manoeuvres, multiple occlusion technique,

multiple linear regression during mechanical ventilation and plethysmography

[32,61,69,74,173]. Table 5.4 shows Crs, Rrs and R a w values obtained from these

investigations in comparison with our values of Crs and Rrs calculated from the 0.5 H z

J

>*~i

>

qou -

300 -

150 -

Ill

impedance data, and the R a w determined by the model fitting.

Table 5.3: Coefficients of variation for respiratory mechanics

Study

Current Study

Turner [47]

Hayden [131]

Consensus Statement [32]

FEV0.5

7.6 ± 0 . 7 %

5.3 ±0.5

5.3 ±1.3

Raw G

11.3 ±0.8% 13.9 ± 1 . 2 %

11-15% (Rrs)

H

9.5 ± 0.7 %

10-13% (Crs)

In all cases the values for Crs reported by previous investigators for a similar study

population were higher than those ascertained in the present study. One possible

explanation for these differences could be the higher average lung volume applied in our

current measurements. Indeed Petak et al [76] reported a significant decrease in the

respiratory elastance (H) when the lung volume decreased from a Ptr of 20 c m H 2 0 to a

level of 10 c m H 2 0 . Further decrease in Ptr, however, caused significant elevations in H.

Therefore, it is more likely that the differences in techniques used in the current study and

those applied previously may explain these differences, since small-amplitude oscillations

result in a significantly smaller Crs, particularly of its chest wall component [4,107].

The Rrs data reported in the present study are comparable with those reported by previous

investigators (Table 5.4). The slight differences can be explained primarily by the higher

lung volume applied in the present study, which causes a marked decrease in the airway

component of Rrs [76].

The values for Raw reported in the literature vary greatly depending on the technique used

(Table 5.4). Lanteri et al. [61] reported values of R a w at 25, 50 and 7 5 % of expired

volume in intubated, mechanically ventilated children using the occlusion technique. The

R a w reported in the present study is higher than that of the R a w at 2 5 % of expired volume

reported by Lanteri et al. (Table 5.4) [61]. The current values for R a w include the upper

airway and nasal resistance while these components were excluded in their study due to the

subtraction of the resistance of the E T T from their estimate of Raw. Plethysmographic R a w

data reported in a consensus statement [32] are 1.5-2 times greater than those obtained in

the present study. Besides the obvious effects of the elevated lung volume on R a w [76], the

112

presence of a flow-dependent component in the plethysmographic R a w due to the higher

respiratory flows may also explain these differences. The mean coefficients of variation

reported in the present study are comparable to those described in the ATS/ERS consensus

statement [32].

Table 5.4: Comparison of respiratory mechanical parameters in healthy infants

Study

Length (cm)

Current Study

Turner [47]

Hanrahan [69]

Lanteri [61]

Masters [74]

FEV0.5 (ml)

60 80

95.0 212.1

121.2 228.6

Raw (cmH20.s/L)

60 80

24.8 16.4

31.0 16.9

Rrs (cmH20.s/L)

60 80

59.7 26.6

60.8 32.8

70.2 19.6

39.2 28.6

Crs (L/cmH20)

60 80

0.0027 0.0080

0.0074 0.0147

0.0070 0.0183

0.0066 0.0143

5.5.2 RVRTC

The number of previous studies detailing forced expiratory parameter values in normal

infants are limited. A small number of studies have reported predictive values for maximal

flow at F R C (VmaxFRC) generated by the R T C [38,44,69]. A previous study from our

laboratory generated predictive values of FEV t for normal infants [47], using a prototype

of the current equipment. The predictive equation for FEV0.5 determined in the present

study is comparable with that published by Turner et al. [47](Table 5.4). The group mean

intra subject C V obtained for FEV0.5 were comparable with those described previously

[47,131](Table5.3).

5.5.3 Factors influencing lung function parameters

Gender was found to have no effect of lung mechanics in the present study, while family

history of asthma was a significant covariate having a detrimental effect on Raw, H and

FEV0.5. Recently Hanrahan and coworkers [69] reported mixed gender effects in a

longitudinal study. Female infants were found to have significantly lower initial levels of Rrs

than males, however the decrease in Rrs with increasing length was slower in female infants.

In contrast, while females had lower initial levels of Crs and a slower decline in Crs than

113

males, these changes were not significant. Indeed, the study of Hanrahan et al. [69]

reported longitudinal data in 541 infants and demonstrated inconsistent gender effects in Crs

compared with Rrs. In addition, Tepper and Reister [38] reported higher forced expiratory

flows in females than males, however the difference was not significant. Overall, the smaller

numbers in this cross-sectional study would account for finding the non significant effect

of gender.

Family history of asthma had a significant detrimental effect on all lung function parameters,

with the exception of Grs and law. This is in marked contrast to the study of Tepper and

Reister [38], w h o found no effect of family history on V m a x F R C or FRC. Young and

coworkers [174] demonstrated an increased level of airway responsiveness to histamine in

infants with a family history of asthma, however baseline lung function levels were not

different between infants with or without a family history of asthma. The Tucson study

[175] reported an effect of maternal history of asthma of infants with late-onset wheezing

(>3) when compared to infants with no wheeze up to the age of 6. In addition, there was

a non significant trend for those children with late-onset wheeze to have lower V m a x F R C

at age < 1 year. The present study only examines lung function up to the age of 2 years and

hence cannot differentiate between those children who will be free of wheeze later in life

and those who may develop late-onset wheeze. It is possible that the infants in the present

study with a family history of asthma and decreased lung function may represent the group

described by Martinez et al [175] as late-onset wheezers.

5.5.4 Dysanaptic growth

Several investigators have reported differing rates of growth of the airways and lung tissues

from infancy through to adolescence [61,176,177]. Mead [176] studied the association

between maximal expiratory flow and vital capacity, while Lanteri et al. [61] examined the

relative changes in R a w and Crs with growth. Hibbert et al [177] studied longitudinal lung

volumes and flows and suggested that as lung function parameters 'tracked' with age then

the rate of growth of the differing lung structures would remain constant relative to each

other. To examine this issue in the present population w e compared the rate of change in

airway and tissue mechanics as previously suggested by Lanteri and Sly [61]. This was

carried out by relating the changes in Ln(Raw) to those in Ln(G) and Ln(H). The slopes of

114

the regression equations (0.47 and 0.49 for G and H, respectively) show the tissue

parameters decreasing faster with length than Raw. The findings from the present study

therefore provide further confirmation of the concept of dysanaptic growth during infancy.

5.5.5 Comparison between FOT and RVRTC

To ascertain the role of the individual oscillatory parameters in predicting FEV0.5 w e

modelled the influence of the airway and parenchymal mechanics on the forced expiratory

volumes within this patient group (n=23). W e found that FEV0.5 was best described by a

model including H and R a w and their squares, the addition of multiplicative interactions

(e.g. HZRaw) did not improve the predictive value of the model. One infant was an outlier

and strongly influenced the model, if this infant is excluded then FEV0.5 was best predicted

by elastance and it's square only (FEV0.5=667.3-0.063(H*Lt)+1.7*10-6(H*Lt)2; R2=78.5%,

pO.0001). This result indicates that, when airway and parenchymal mechanics are used to

predict forced expiratory volumes, airway caliber and elastic properties of the parenchyma

are crucial factors, while inertance and tissue damping are secondary contributors. The

small numbers used in the present study do not allow definite conclusions to be reached,

studies with larger numbers are required to minimize the effect of outliers.

5.5.6 Conclusions

The present study describes changes in oscillatory and forced expiratory mechanics with

growth in normal infants using the techniques of low-frequency forced oscillation and

R V R T C in the first two years of life. This information has been used to establish the

relationship between values of respiratory mechanics and length in sedated normal infants.

Evidence supporting the concept of dysanaptic growth between the airways and pulmonary

parenchyma is detailed with the respiratory tissues found to decrease more rapidly with

growth than the airways. The relationships between airway and parenchymal mechanics and

measures of forced expiratory volume were also explored. FEV0.5 was strongly predicted

by elastance and airway resistance confirming that airway caliber and elastance have clinical

implications within both forced expiratory and oscillatory mechanics.

115

CHAPTER 6: RESPIRATORY SYSTEM IMPEDANCE IN WHEEZY INFANTS

6.1 Summary

The low-frequency forced oscillation technique (FOT) was used to measure the mechanical

properties of the airways and respiratory tissues in 22 infants with a history of recurrent

wheeze. All infants were asymptomatic at the time of testing and had been free of upper

respiratory tract infections for at least 4 weeks. The Hering-Breuer reflex was used to

induce a pause in breathing, allowing low-frequency respiratory impedance spectra (Zrs)

between 0.5 and 21 H z to be obtained. A model containing an airway (resistance; R a w and

inertance; law) and a constant-phase tissue (tissue damping; G and tissue elastance; H )

compartment was fitted to the respiratory impedance spectra. Standardized variants (Z

scores) were calculated to compare the infants to a normal population. Infants with a

history of recurrent wheeze had increased respiratory tissue (Z scores: 0.54±0.23 and

0.22±0.01; p<0.05 and p<0.01, for G and H, respectively), but not airway mechanics ( 0.18

±0.23) when compared to regression equations derived from a normal infant population.

Infants under the age of 1 were not significantly different from normals in any oscillatory

parameter (-0.29±1.36, 0.02±0.24 and 0.15±0.23, for R, G and H respectively) while those

infants > 1 year had elevated airway and respiratory tissue mechanics (0.57±0.18,

0.97±0.32 and 1.31±0.28; p<0.01 for R and G, and pO.OOl forH, respectively). These

results indicate that the low-frequency F O T is able to discriminate recurrently wheezy

infants from normal infants as a group, that infants with recurrent wheeze that persisted

beyond 1 year may represent those infants w h o go on to develop asthma and that the

mechanics of both the airways and respiratory tissues are abnormal in infantile wheeze.

116

6.2 Introduction

In early infancy rapid changes occur in the physiological, immunological and anatomical

development of the respiratory system. Until recently the available knowledge regarding the

normal physiological development of the lung was limited and as a result, the understanding

of the development of disease and responses to treatments has suffered. Wheeze in infants

may result from a number of anatomical, physiological or immunological processes not all

of which may contribute to asthma in later life. Recent epidemiological data has identified

3 distinct groups in the development of respiratory disease [175]. Those infants w h o

wheeze, with a lower respiratory tract infection (LRI) in the first 3 years of life but do not

wheeze or have asthma at 6 years of age (early wheeze). Late-onset wheezers, including

those infants with no wheeze early in life, but w h o develop wheeze between the ages of 3

and 6 years of age, and those infants who wheeze early in life and still have wheeze by the

age of 6 (persistent wheeze). Different risk factors have been associate with each of these

groups. Those infants with transient wheeze had diminished premorbid V m a x F R C and may

represent those infants with congenitally smaller airways, hence predisposing them to LRI

and wheeze in early life. [175]. Infants with persistent wheeze had initial lung function

values similar to those infants without wheeze, but decreased V m a x F R C at age six,

suggesting other mechanisms for infantile wheeze than small airways. These infants also had

more frequent symptoms and were more likely to have had rhinitis, eczema, maternal

asthma and elevated serum IgE at 9 months than the group of transient wheezers [175].

While it is now possible to surmise which group an infant may fall into, w e are currently

unable to classify with any degree of certainty those infants who may go on to develop

asthma and those infants w h o may grow out of their wheeze.

Recent advances in lung function measurements have allowed a number of investigations

to be carried out in normal infants and those infants with acute and recurrent wheeze. In

older children and adults the incidence of asthma is associated with decreased lung function,

increased responsiveness to bronchoactive agents and elevated levels of serum IgE. Yet

while decreased levels of premorbid lung function have been observed in infants with early-

onset wheeze [175,178,179], increased bronchial responsiveness has not been shown to be

correlated to a history of wheeze in infancy [178]. A number of lung function parameters

have shown differences between those infants with a history of wheeze and normal infants.

117

These include length and F R C corrected V m a x F R C , but not F R C [179] and specific airway

conductance (sGaw) [180]. A further study comparing the single-breath (SBT) and

plethysmographic techniques have shown differences between Rrs and Crs [181]. A study

combining the single-breath method and the forced oscillation technique has not shown

differences between normal infants and those with a history of airway obstruction (including

bronchiolitis, chronic lung disease and cystic fibrosis) in any measured parameter [182]. The

techniques described above are limited as they are unable to provide a separate assessment

of the airway and lung tissue mechanics and cannot definitively identify the site of changes

in lung function. The low-frequency F O T allows the measurement of the Zrs below the

spontaneous breathing frequency and thus the simultaneous determination of airway and

tissue mechanics. W e hypothesized that by examining the airway and lung tissue mechanics

in infants with a history of recurrent wheeze and comparing the results with previously

determined regression equations for normal infants (See Chapter 5), that the mechanical

properties of both the airways and tissues would contribute to respiratory disease in infancy.

6.3 Methods

6.3.1 Subjects

Twenty-two infants with a history of recurrent or persistent wheeze were examined, all

infants were asymptomatic and had been free of symptoms for a period of at least 4 weeks

prior to testing. Recurrent wheeze was define as 3 or more episodes of wheeze in the past

12 months of life, persistent wheeze was a period of consistent wheezing exceeding 1

month. Anthropometric data on the infants are outlined in Table 6.1. The infants were

sedated with an oral dose of choral hydrate (70-100 mg/kg) and were laid in the supine

position with the head supported in the midline and the neck slightly extended. The

experimental protocol was approved by the H u m a n Ethics Committee of the Princess




See Section 2.2.2.1 for complete details.

The input impedance of the respiratory system (Zrs) was measured as described previously

[7]. Briefly the infants lungs are raised to a transrespiratory pressure of 20 c m H 2 0 three

118

times and the airway occluded inducing the Hering-Breuer reflex. During the resulting

pause in breathing a forcing signal between 0.5 and 20 H z was applied to the airway

opening, via a soft-rimmed face mask by a loudspeaker. A lumped parameter model of the

respiratory system was fitted to the individual Zrs data [26]. The model contained a

frequency independent airway resistance (Raw) and inertance (law) and a tissue

compartment including tissue damping (G) and elastance (H) terms that alter in phase with

each other.

6.3.3 Study Protocol and Analysis

Baseline Zrs data were collected on those infants with a history of recurrent or persistent

wheeze, previously enrolled in other protocols. Individual airway and respiratory tissue

mechanics were obtained and averaged. Baseline mechanics were then compared to the

regression equations derived in Chapter 5. Standard variants, Z scores, were calculated as

the difference between the measured and predicted pulmonary function, divided by the

residual standard deviation derived from the regression equations. A t-test was carried out

to test the null hypothesis of the z scores. As the definition of recurrent wheeze was 3 or

more episodes of wheeze in the last 12 months of life and a number of infants were under

the age of 1 year (n=10) a separate analysis comparing infants < 1 year with those infants

>1 year of age was carried out using a two-tailed t-test. Significance was accepted at the

0.05 level.

119

Table 6.1: Anthropometric data and lung function data

Patient

1

2

3

4

5

6

7

8

9

10

11

12

13

14

15

16

17

18

19

20

21

22

Mean(SEM)

Age

0.48

0.50

0.90

0.66

0.82

0.95

1.20

1.39

0.99

0.90

0.84

1.52

1.64

1.41

1.53

0.99

1.31

1.92

1.61

1.89

2.06

1.63

1.23(0.1)

Length

67.0

68.5

70.0

72.0

74.0

74.0

74.0

75.0

75.8

77.5

78.0

79.0

80.0

80.5

81.0

82.0

82.0

82.0

87.0

87.0

87.5

89.0

78.3(1.3)

Raw

21.5

28.5

13.3

22.0

11.7

30.8

22.2

18.4

12.9

13.7

14.3

23.5

17.6

22.7

23.4

20.9

18.9

14.1

13.3

14.0

16.7

12.0

18.5(1.2)

law

0.12

0.09

0.09

0.18

0.07

0.24

0.08

0.20

0.08

0.05

0.12

0.25

0.10

0.21

0.19

0.09

0.17

0.04

0.10

0.07

0.11

0.03

0.12(0.01)

G

69.7

62.0

41.2

43.7

32.3

40.8

26.3

55.1

25.7

24.1

32.4

41.1

28.5

35.0

32.4

30.6

30.6

29.3

28.7

18.4

21.9

37.4

35.8(2.7)

H

222.4

214.1

172.3

169.1

153.9

152.8

132.6

196.7

116.1

131.3

146.4

134.3

117.3

123.7

150.2

141.2

124.2

128.0

105.0

72.9

93.5

114.4

141.5(7.8)

T|

0.32

0.29

0.24

0.26

0.21

0.27

0.20

0.28

0.22

0.18

0.22

0.31

0.24

0.28

0.23

0.22

0.25

0.23

0.27

0.25

0.22

0.33

0.25(0.01)

Table 6.1: Anthropometric and baseline lung function data. Age (years); Length (cm); Raw (cmH20.s/L); law

(cmH20.s2/L); G (cmH20/L) and H (cmH20/L)

120

6.4 Results

Figure 6.1 shows the mean lung function measurements for each infant. Infants below the

age of 1 are scattered evenly around the regression line, while those infants older than 1

year of age tend to skew above the regression line indicating a worsening of lung function

when compared to healthy infants. The testing of the null hypothesis of the Z scores

revealed that only the respiratory tissue mechanical parameters were significantly different

(Table 6.2). There were no significant differences in recurrently wheezy infants below 1

year of age when to compared to healthy infants, while significant differences were noted

in all lung function parameters in those infants older than 1 year (Table 6.2).

Table 6.2: Z score data for wheezy infants

Z score

(mean±SEM)

R a w

G

H

<1 year

-0.29(1.36)

0.02 (0.24)

0.15(0.23)

> 1 year

0.57 (0.18) §

0.97 (0.32) §

1.31(0.28)U

Total population

0.18(0.23)

0.54 (0.23) *

0.22 (0.01) §

Table 6.2: * ; p< 0.05, § ; p< 0.01 and ̂ ; pO.OOl.

6.5 Discussion

The data in the present study indicate that recurrently wheezy infants have abnormal airway

and parenchymal mechanics when compared to a normal population. Previously published

data have demonstrated altered lung function in wheezy infants when compared to normal

infants [178,181]. In addition Tager et al [179] reported decreased length- and F R C

corrected-VmaxFRC values, while Dezateux et al. [180] reported decreased specific

conductance (sGaw) in infants, tested at an early age, who then subsequently develop

wheeze. Dundas and coworkers [181] demonstrated increased Rrs and decreased Crs, in

wheezy compared to normal infants, while Stick etal. [178] reported decreased V m a x F R C

values. In contrast a number of lung function parameters have been shown to be unaltered

in wheezy infants including; tidal volume, Crs and respiratory system conductance [178],

F R C [181] and Raw, thoracic gas volume (TGV) and specific R a w [183]. Longitudinal

121

studies have demonstrated lower baseline lung function prior to wheeze when compared to

infants w h o do not subsequently develop wheeze [175,179,180]. While methodological

differences may account for some of the reported differences in the previous studies, it is

not clear from the previous reports which infants with wheeze have diminished lung

function. The ability to detect those infants w h o wheeze due to anatomically, small for age

airways may increase the sensitivity of infants pulmonary lung function tests to accurately

identify wheezy infants w h o will go on to develop asthma.

The current study demonstrated abnormal airway and respiratory tissue mechanics in

wheezy infants older than 1 year of age, but not in infants younger than this. In infants less

than 1 year of age, this may represent a single winter season and be the result of repeated

acute viral wheeze, rather than ongoing chronic wheeze. The diagnosis of recurrent wheeze

in the older infants may better represent chronic wheeze and more accurately identify those

infants w h o may go on to develop asthma. Respiratory system resistance measurements

include both the upper- and lower airways and the tissue properties of the lung and chest

wall. Specific airway conductance can provide a measure of airway calibre corrected for

lung volume and is thought to be a sensitive index of airway function [184]. The advantage

of the low-frequency F O T is it's ability to simultaneously provide separate information on

the airways and respiratory tissues. Our estimates of R a w will include the resistance of the

nose and the extra-thoracic airways. Data reported elsewhere in this thesis (Chapter 4)

demonstrates that the nasal impedance contributes approximately 4 5 % to R a w and that this

contribution is constant with length. The presence of an upper respiratory tract infection

(URI) is likely to influence upper airway mechanics and infants were excluded from the

analysis if an U R I had occurred in the previous 4 weeks. The addition of a nasal catheter

would increase the sensitivity of the low-frequency F O T to detect changes in peripheral

airway function and may help identify those infants with anatomically small for age airways.

122

^40

1

o

20 w rt

.5

Length (cm) 80 90

Figure 6.1: Mean individual data for Raw, G and H in wheezy infants. The hollow points are those infants < 1

year of age, while the solid points represent those infants > 1 year. The solid line represent the regression

equation presented in Chapter 5. The variation in data representing standard variants of 1 (long dash) and 2

(short dash) are also shown.

The contribution of the chest wall to respiratory system resistance and elastance, increases

with age as the chest wall becomes stiffer (See Chapter 4 for full details). The estimates of

G and H in the present study will include the contribution of the chest wall. Changes in

123

chest wall motion have been reported with pulmonary disease, with the chest wall and

abdominal movements being asynchronous, this asynchrony then decreases with improved

lung mechanics [32]. However it is unlikely that chest wall damping and elastance would

be abnormal in recurrently wheezy infants. Wheeze in infancy and subsequently asthma is

characterized by abnormal expiratory flows and implies flow limitation [185]. A s lung

volume decreases the flow limiting segment shifts peripherally and hence flows at F R C are

more likely to be influenced by the mechanical properties of the small airways and the

mechanical interdependence of the airways and parenchyma [186]. Changes in the

mechanical properties of both the airways and respiratory tissues, that may occur with

wheeze in infancy have not been systematically addressed. The majority of infant lung

function studies utilize forced expiratory flows and/or volumes to assess lung mechanics in

spontaneously breathing infants. While airway obstruction will effect expiratory flow

limitation, the mechanical properties of the parenchyma will also play a role. Direct

conclusions as to any alterations in lung tissue mechanical properties cannot be made from

measurements of forced expiratory flows. The parenchyma may effect the expiratory flow

limitation in a number of ways; general lung elastic recoil, localised (peribronchial) recoil

and airway-parenchymal interdependence [187]. A change in any of these factors will

increase or decrease the effect which the pulmonary parenchyma has on the flow limitation,

and hence wheeze. The lung elastic recoil forces determine the driving pressure within the

lung, while the peribronchial recoil forces and airway-parenchymal interdependence will

apply distending pressure to the outer surface of the airways and hence influence airway

calibre to some extent. The elastic recoil of the lung is created by the connective tissues of

the lung and by the gas-liquid interface. Increased dynamic elastance has been reported in

asthmatic subjects during quiet breathing [188], in addition elastic recoil is reduced in adult

asthmatics [189]. In the small intraparenchymal airways the connective tissues of the

parenchyma are directly connected to the connective tissues of the airways, this connection

will alter airway-parenchymal function if one component is changing disproportionally

compared to the other (e.g. with airway constriction) [187]. Airway-parenchymal

interactive forces play a significant role at high lung volumes where elastic forces acting on

the connective tissues are greatest and may be changed by localised changes to the

parenchyma (e.g. edema or inflammation), while elastic recoil and peribronchial recoil will

be susceptible to increased tensions in the force-bearing elements surrounding the airways

124

[187]. Kraft et al [190] have demonstrated while inflammation is present in the proximal

airways and alveolar tissues in asthmatics, that it was alveolar eosinophil and macrophage

infiltration, rather than airway infiltration, that was responsible for nocturnal decreases in

lung function associated with episodes of nocturnal asthma. This infiltration may be acting

to alter the localised and general recoil forces within the lung parenchyma, hence altering

the expiratory flow limitation. The abnormally high levels of tissue damping and elastance

seen in the infants with ongoing wheeze at 12 months may represent the beginnings of

chronic tissue inflammation and may represent those infants w h o will go on to develop

asthma later in life, confirming that the pulmonary parenchyma play an important role in

respiratory disease and that conditions such as asthma may reflect alterations in both the

airways and respiratory tissues.

6.5.2 Conclusions

W e have shown that the low-frequency forced oscillation technique determined at a

transrespiratory pressure of 2 0 c m H 2 O is able to detect abnormal lung function in infants

with a history of recurrent wheeze. The mechanical properties of the airways and

respiratory tissues were found to be abnormal in infants older than a year, but not younger

than this age. Respiratory tissue mechanics were also abnormal for the population as a

whole. The finding of this study indicates that the airways and pulmonary tissues play a role

in infantile wheeze. Further studies investigating the role of passive smoke exposure, family

history of asthma and the possible effect of anatomically small for age airways are needed

to better understand the importance of factors, such as the physiological, anatomical and

immunological development of the respiratory system, in the first years of life to the

subsequent development of asthma.

125

CHAPTER 7: METHACHOLINE

RESPONSIVENESS IN INFANTS

7.1 Summary

The usefulness of the low-frequency forced oscillation technique (FOT) to detect the

response to inhaled methacholine (Mch) was examined in 7 infants with no history of

respiratory disease, 8 with recurrent wheeze and 2 with recurrent cough without wheeze,

all infants were asymptomatic at the time of testing. The respiratory system impedance (Zrs)

between 0.5 and 21 H z was determined during a pause in breathing induced by the Hering-

Breuer reflex. Zrs were fitted to a model containing a frequency independent airway

resistance (Raw) and inertance (Taw) and a constant-phase tissue damping (G) and elastance

(H). Hysteresivity (n.) was calculated (G/H). The raised volume rapid thoracic compression

technique (RVRTC) was used to generate forced expiratory volume in 0.5s (FEVo.s). Lung

function was determined at baseline and following inhaled Mch, in doubling doses from 0.25

to 16 mg/mL, until the maximal dose was reached or a drop of 1 5 % in FEV0.5 was achieved

(PC15FEV0.5). The concentration required to cause an increase in R a w twice that of the

baseline variability (1.1±0.4 mg/mL Men; mean±SEM; TCRaw) was significantly lower than

PC15FEV0.5 (4.7±1.7 mg/mL Mch). At the T C R a w a response in Raw, law, G and rj, but

not H was detected (34.2±6.3%, 67.2±14.1%, 33±3.7%, 31.9±12.9% and -2.3±3.7%,

respectively). The changes required to cause an increase in R a w of 3 0 % and a decrease of

1 5 % in FEV0.5 are approximately equal multiples of the respective baseline variabilities

(-2.5-3 time the coefficient of variation). Furthermore it was found that PC3oRaw and

PC15FEV0.5 were not significantly different. The low-frequency F O T is able to detect

changes in airway and respiratory tissue mechanics following inhaled Mch. A n increase of

3 0 % in R a w would be a suitable outcome variable for further studies.

126

7.2 Introduction

Responsiveness to an acmiinistered drug or compound remains the most useful physiologic

test in the assessment of asthma in adults and older children. A number of investigators have

adapted the protocols used in adults and older children for use in infants [41,131,174].

Airway responsiveness has been shown to be increased in infants with a family history of

asthma [174], in infants with cystic fibrosis [191] and to correlate to baseline lung function

(VmaxFRC%pred) in normal infants and those infants with cystic fibrosis [191]. While the

presence of reversible hyperresponsiveness is well known in asthmatic adults, the role of

airway responsiveness in whee-zy infants is unclear. Stick et al [178] demonstrated that

while infants with a history of recurrent wheeze had significantly lower baseline lung

function when compared to normals, that the responsiveness of the two groups to inhaled

Wstamine were equal, in contrast Clark et al[192] reported an increased responsiveness in

female neonates w h o subsequently went on to develop lower respiratory tract infection

(LRI) in the first year of life.

Several techniques have been used to assess the bronchial responsiveness in infants, these

include the raised volume rapid thoracic compression technique ( R V R T C ) [131], the rapid

thoracic compression technique (RTC) [43], the forced deflation technique [193] and the

high-frequency forced oscillation technique [121]. These techniques are unable to provide

separate information on the behaviour of the respiratory tissues. A number of investigations

in animals have demonstrated that the pulmonary tissues do indeed respond to inhaled

bronchoconstrictors [5,139,194-196]. The low-frequency forced oscillation technique

(FOT) allows the mechanical behaviour of the respiratory system airways and tissues to be

characterized. The impedance spectra of the respiratory system (Zrs) has been well

characterized between ~2 H z and 48 H z [102,122,182,197] and at higher frequencies (>48

Hz) [106,121]. Recently Sly et al. [7] demonstrated that by utilizing the Hering-Breuer

reflex, a pause in breathing could be induced that allowed the reliable measurement of Zrs

at low frequencies (<2 Hz). Using this method Hayden and coworkers [123] demonstrated

a significant fall in airway resistance (Raw) and a decrease in tissue damping (G) following

the administration of a bronchodilator, concluding that low-frequency F O T was suitable for

studying the bronchodilator response in the airways and respiratory tissues in infants.

127

The aim of this study was to assess the effectiveness of the low-frequency F O T to

characterize the response to inhaled methacholine (Mch) in sedated infants. W e compared

the response to inhaled M c h in infants using the R V R T C and the low-frequency forced

oscillation techniques using an adaptation of the methodology used by Hayden et al. [131].

7.3 Methods

7.3.1 Subjects

Seventeen infants were enrolled in the study. Eight infants had a history of 3 or more

episodes of wheeze in the last 12 months of life (recurrent wheeze), two infants had 3 or

more episodes of cough without wheeze (recurrent cough), the remaining 7 infants had no

history of respiratory disease (normal). All infants had been free of symptoms or illness for

a period of at least 4 weeks. Anthropometric data are presented in Table 7.1. The infants

were recruited from the general public. Parents gave written informed consents and were

generally present during the study. The study was approved by the relevant Institutional

Human Ethics Committee. Infants were sedated with an oral dose of choral hydrate (70-100

mg/kg), heart rate (HR) and oxygen saturation (Sa02) were monitored throughout the

study. Infants were studied in the supine position with the head supported in the midline and

slightly extended.


7.3.2.1 Raised volume rapid thoracic compression technique


Briefly, the method of Hayden et al. [131] was used to determine forced expiratory volume

at 0.5s (FEV0.5). This technique uses a pump to raise the infant's lung volume above the

tidal range. The pump draws room air through a pneumotachograph and into the infant,

using a series of solenoid valves the infants were inflated three times to a transrespiratory

pressure of 20 cmH 20 with passive deflations between each inflation. A jacket is connected

to a positive pressure reservoir and a compression force was applied to the thorax and

abdomen after the third and final inflation. Forced expiratory flow was determined,

integrated and volume-time curves were produced. FEV0.5 was calculated from the forced

expiratory volume-time curves.

128

Table 7.1 Anthropometric data

Patient

1

2

3

4

5

6

7

8

9

10

11

12

13

14

15

16

17

Age (weeks)

19.6

48.7

47.0

26.1

42.7

24.1

51.6

67.9

100.3

92.7

84.9

66.7

68.1

86.0

97.9

102.6

99.3

Height (cm)

67.5

70.0

70.0

70.0

74.0

74.0

75.8

80.0

80.8

81.0

81.7

81.8

82.0

82.4

82.6

85.0

86.0

Weight (kg)

8.3

7.9

8.2

8.9

9.5

8.9

11.6

10.6

11.6

11.2

11.2

13.5

10.8

11.3

12.6

14.0

12.0

Gender

M

F

F

F

M

F

M

F

F

M

F

M

M

F

F

F

M

Classification

Normal

Cough

Wheeze

Normal

Wheeze

Wheeze

Wheeze

Normal

wheeze

Normal

Normal

Wheeze

Wheeze

Normal

Cough

Normal

Wheeze

7.3.2.2 Low-frequency forced oscillation technique

See Section 2.2.2.1 for full details.

Respiratory input impedance spectra (Zrs) were measured by using the low-frequency

forced oscillation technique as described by Sly et al [7]. Three inflations were applied to

a circuit containing the infant, the loudspeaker and a pressure reference chamber. Following

the third inflation, the airway was occluded at a transrespiratory pressure (Ptr) of 20 cmH 20

and in the resulting pause in breathing, the oscillatory signal was driven into the infant's

respiratory system by the loudspeaker. Following each measurement the mask was lifted

and the infant allowed to breathe spontaneously. A linear model of the respiratory system

[26] was fitted to the individual respiratory impedance spectra in the 0.5 -15 Hz range (see

Section 2.4).

129

7.3.3 Study protocol

The protocol used was a modified version of Hayden and coworkers [131]. Lung function

was assessed firstly using forced oscillation and then R V R T C techniques at baseline and

following inhaled nebulisations of saline and Mch. Nebulisations were given for 1 min, of

tidal breathing at 5 min intervals. M c h nebulisations started at a concentration of 0.25

mg/mL and increased in doubling doses until a positive response was recorded or to a

maximum of 16 mg/mL (0.25, 0.5, 1, 2, 4, 8,16 mg/mL). The test was discontinued if the

Sa0 2 dropped below 8 5 % at any time throughout the test. Each set of raised volume and

forced oscillation data was analyzed to produce outcomes for FEV0.5 and oscillatory

parameters (Raw, law, G, H and r|), respectively. Mean values of FEV0.5 were calculated

for each dose of Men, with a fall in FEV0.5 of 1 5 % from control considered a positive

response to M c h [131]. 200 u.g of salbutamol was administered, via a pressurized metered

dose inhaler (MDI) and small-volume metal spacer (Nebuchamber, Astra Pharmaceuticals,

Sweden) following the completed M c h protocol.

7.3.4 Analysis

Baseline and control measurements of FEV0.5 and oscillatory parameters were averaged

(mean±SD). As detailed above, a 1 5 % fall in FEV0.5 from the control measurement

(PCisFEVo.5)was considered a positive response to M c h and the test was stopped.

Respiratory impedance spectra (Zrs) were examined and the peak measurement for each

M c h dose analyzed (see Section 2.2) and reported. Oscillatory parameters were analyzed

as the percentage increase from control. The infant was considered to have had a positive

response to Mch, as determined by the low-frequency FOT, if the R a w percentage increase

from control exceeded 2 times the coefficient of variation (i.e. the threshold concentration)

and remained above that level for the remainder of the study (TCRaw). The relative

sensitivities of the two tests were compared using a Wilcoxon rank test. Differences in

baseline lung function and M c h responses in the patient subgroups were also carried out

using one-way A N O V A . Significance was accepted at the p<0.05 level.

7.4 Results

7.4.1 Forced expiration

130

The mean baseline values for FEV0.5 were 201.9 ± 12.4 m L (mean ± S E M ) (Table 7.2).

There were no differences between the subgroups of patients. Fourteen of the 17 infants

had a 1 5 % fall in FEV0.5 before waking up. Of the remaining 3 infants, 1 woke due to cough

and/or wheeze during nebulisation, with the remainder waking spontaneously without

obvious cause. Table 7.3 shows the concentration of Mch causing a 15% decrease in FEV0.5

(PC15 FEVo.5).

•7.4.2 Low-frequency forced oscillation technique

Figure 7.1 shows the dose response curves for patient #14. The mean control data includes

an error bar indicating the increase required to exceed 2 times the COV. Individual peak

measurements following inhaled Mch are also shown. Baseline values for the oscillatory

parameters are shown in Table 7.2. The z-score values for the lung function parameters

were calculated using the predictive equations reported in Chapter 5. There were no

differences in baseline data or z-score data between any subgroup of the infants. Z-score

data for H was significantly different from the normal population described in Chapter 5 (§;

p<0.01). The fitting error was calculated under baseline and maximally constrictive

conditions. The model was found to fit the Zrs well following inhaled Men, with no

significant differences seen in fitting error from control conditions (7.6+ 0.3% and

8.1+0.4%, control and Men, respectively).

Table 7.2: Baseline lung function data

Lung Function Parameter

FEVo.5 (mL)

Raw (cmH20.s/L)

law (cmH20.s2/L)

G (cmH20/L)

H (cmH20/L)

n

Mean (SEM)

201.9 (12.4)

18.3 (1.2)

0.11(0.01)

29.4 (3.0)

145.8 (10.2)

0.21 (0.01)

Z score (SEM)

-0.28 (0.23)

0.10(0.25)

-0.41 (0.22)

0.77 (0.23)§

COV(SEM)%

3.8(0.7)

8.3 (1.3)

11.5(1.7)

14.4 (1.9)

8.7(1.4)

18.1(2.1)

131

100

a 75

I 1 0 s

50

25

0-

Airway Resistance 50

25

0 -

Airway Inertance

i — i — i r

I 1

150

125

100

75

50-

25

Tissue Damping

16

20

10-

-10 -

Tissue Elastance

u.

T — I — i r-

SI 2 4 16 M c h (mg/mL) Mch (mg/mL)

Figure 7.1 Representative cumulative dose response curve to inhaled Mch (patient 14#). The percentage

increase from saline are shown. The error bar indicates the increase needed to exceed twice the baseline

variability. A positive response to Raw, law and G can be seen, no reponses in tissue elastance was elicited.

Three of the 17 infants woke before completion of the test (administration of 16 mg/mL

M c h or 1 5 % decrease in FEV0.5), of the 3, one had already had a positive response to M c h

in airway resistance. In the 14 infants w h o remained asleep for the duration of the study,

9 had an increase in R a w greater than 2 times the C O V . In 5 of the 17 infants the apnoeic

pause in breathing was insufficient to allow the reliable measurement of Zrs. In all cases this

occurred at the M c h concentration that the FEV0.5 decreased by 15%, this represents an

ongoing difficulty with the cumulative dose response challenge test as carried out in this

investigation. Table 7.3 shows the threshold concentration of M c h causing an increase in

132

R a w greater than 2 times the C O V (TCRaw) and the corresponding percentage increases

from control for the remaining oscillation parameters at the TCRaw, those data marked

with * indicates a positive response in that parameter to M c h at the TCRaw.

A Wilcoxon signed rank sum test was carried out to determine the differences between the

M c h concentrations required to elicit a positive response in FEV0.5 and Raw. The threshold

concentration for R a w was significantly lower than the PC15FEV0.5 (Table 7.3; p<0.05). N o

differences in PC15FEV0.5 or T C R a w were seen between normal infants and those infants

with wheeze or cough.

Table 7.3: Provocation and lung function data

Patient

1

2

3

4

5

6

7

8

9

10

11

12

13

14

15

16

17

Mean(SEM)

PCi5FEVo.5

1.0

1.0

1.0

16.0

0.5

4.0

1.0

0.25

Aw

16.0

Aw

Aw

2.0

16.0

4.0

2.0

1.0

4.7(1.7)

TCRaw

Br

0.25

1.0

2.0

0.5

1.0

0.5

0.25

Aw

Br

1.0

Aw

Br

4.0

Br

0.25

Br

1.1(0.4)

%Raw

—

11.22

56.82

34.17

44.53

20.49

22.42

76.09

—

~

27.16

—

~

32.75

~

16.08

~

34.2(6.3)

% law

—

15.61

95.76*

55.99*

48.95*

68.36*

105.60*

164.47*

--

—

18.98

—

~

44.86*

«

59.05*

~

67.8(14.1)

%G

«

10.08*

18.03

5.21

20.86*

23.81*

116.29*

62.57*

--

—

8.22*

--

—

26.75

—

38.60*

—

33.0(0.7)

%H

~

2.67

-19.95

-8.93

9.77

-0.09

-7.77

-17.17

—

~

18.66*

~

—

-1.76

--

1.75

—

-2.3(3.7)

%r]

—

8.70

46.30

13.65

23.70

22.60

139.70

13.96

—

—

-9.31

—

—

24.10

—

36.04

—

31.9(12.9)

Table 7.3: The provocation concentration causing a 15% decrease in FEV0.5 (PQ5FEV0.5) and an increase in

R a w at least twice the C O V (TCRaw). 'Br1: those infants in w h o m a pause in breathing could not be induced,

Aw': those infants w h o m awoke at that concentration of Mch. The percentage increase from baseline for the

oscillatory mechanical parameters at T C R a w are also shown * indicates a response at the TCRaw.

133

7.5 Discussion

Current methodologies to test the responsiveness to bronchoconstrictors in infants,

commonly assesses changes in V m a x F R C or FEVi [185] and are unable to determine the

site of changes in lung function. The low-frequency forced oscillation technique has been

shown to reliably measure the airway and lung tissue mechanics in infants [7]. The aim of

the current study was to assess the ability of the low-frequency F O T to detect changes in

airway and/or pulmonary tissue mechanics following the administration of inhaled Mch.

The results demonstrate that the technique is able to track changes in lung function due to

inhaled M c h and that the threshold concentration of M c h required to cause an increase in

Raw, twice that of the intrasubject C O V , can be determined. However in 5 of the 17

infants (29.4%) a response in R a w could not be determined, as a pause in breathing of

sufficient length was not able to be induced. In all 5 of these cases this occurred at the

concentration of M c h at which the response in FEV0.5 was subsequently determined.

7.5.1 Comparison between low-frequency FOT and RVRTC

While the success rate of the low-frequency F O T was lower than that of the R V R T C

technique, the median dose required to reach the threshold concentration for R a w was

significantly lower than that required to reach the PC15FEV0.5 (0.75 vs. 1.5 mg/mL Mch;

p<0.05). This result may be due to the relative changes required to reach the T C R a w and

PC15FEV0.5. The intrasubject variability in FEV0.5 in the present study was 3.8±0.7%, thus

a decrease in FEV0.5 of 1 5 % represents a change approximately 4 times that of the baseline

variability. This variability is lower than that reported previously in this thesis (7.6±0.7 %;

See Chapter 5, Table 5.3), and in previous studies (-5%) [47,131], thus a decrease of 1 5 %

would represent a change of 2-3 times the previously reported baseline variability in FEV0.5.

The threshold concentration, is by definition, the concentration required to cause a change

twice that of the baseline variability, thus the differences between T C R a w and PC15FEV0.5

are not unexpected. Using the mean intrasubject variability in R a w reported previously in

this thesis (11.3±0.8 %; See Chapter 5, Table 5.3) a change of 23-34% would be required

to represent changes of 2-3 times that of the baseline variability in Raw. A provocation

concentration leading to a 3 0 % increase in R a w (PC3oRaw) would denote a change of 2.7

times that of the baseline variability and would represent a similar proportional change as

a decrease in FEV0.5 of 15%. Indeed if w e compare the PC3oRaw and PC15FEV0.5 in each

134

infant there are no significant differences (1.5; 0.5-2.0 (median; 25-75% confidence

intervals) vs. 1.5; 1.0-4.0 for PC 3 0 R aw and PC15FEV0.5, respectively).

7.5.2 Mch responsiveness in normal compared to wheezy infants

N o differences between subgroups of patients either in baseline lung function or bronchial

responsiveness was seen with either technique. This result concurs with previous studies

[131,178]. The reasons for the lack of increased responsiveness in infants with a history of

recurrent wheeze are not yet completely clear and could be due to a number of factors. A

Perth based longitudinal study demonstrated increased responsiveness to histamine in

infants, at 4 weeks of age, with a family history of asthma or exposure to environmental

tobacco smoke [174]. The Tucson study [175] demonstrated an increased risk of asthma

at age 6 in those infants with both a family history of asthma and a history of wheeze in the

first years of life. However infants may wheeze due to viral illness and/or due to

anatomically small airways and thus infantile wheeze may not be related to subsequent

development of asthma in childhood [175]. The numbers in the present study are too small

to draw conclusions and further studies with larger numbers, taking into account family

history of asthma and passive tobacco smoke exposure would be required to accurately

determine any differences between infants with a history of wheeze and normals.

7.5.3 Airway versus parenchymal responsiveness

The present study demonstrates a response to inhaled M c h in both the airways and

pulmonary tissues in infants, regardless of respiratory history. The mean increase in law at

T C R a w was 67.8 ± 14.4%, indicating a constriction in the nasal- and/or upper airways. The

inertance of the nasal passages contributes about 7 2 % of the total Irs (See Chapter 4),

making this parameter highly sensitive to changes in the properties of the extrathoracic

airways. Tepper and Steffan [170] demonstrated a decrease in peak expiratory flow (PEF),

but not V m a x F R C following the administration of nasally instilled Mch, concluding that

the topical application of M c h into the nose, resulted in changes in nasal- but not lower

airway resistance. In the same study inhaled M c h resulted in decreases in P E F and

V m a x F R C and was attributed to constriction of the lower airways. Using a catheter-tipped

manometer lodged in the right lower lobe bronchus, Yanai and coworkers [198] have

quantified the relative proportion of the peripheral airway resistance to total pulmonary

135

resistance in adults. Using this technique in asymptomatic asthmatic adults Ohrui et al

[199] demonstrated that inhaled M c h acted on both the central and peripheral airways.

Similar results are reported in the present study, with the increases in R a w and law

suggesting that inhaled M c h in infants acts both in the central and peripheral airways.

Conflicting responses within the pulmonary tissues were seen, with an increase in tissue

damping (33.0±0.7%) and ri (31.9±12.8%), but not in H (-2.3*3.7%) at TCRaw. If the lung

tissue responses are further examined, 9 infants reached the threshold concentration for G

without recording a response in tissue elastance, in addition a further 3 infants exhibited a

response in H without reaching a response in G. The contribution of the chest wall on Zrs

will result in decreased sensitivity of the low-frequency F O T to detect changes in the

viscoelastic properties of the pulmonary tissues (Chapter 4). However the responses

detected in G in a majority of infants makes this explanation for a lack of response in H

unlikely. Lutchen et al. [28] has speculated that peripheral inhomogeneity of airways due

to constriction may cause increases in tissue damping unrelated to actual tissue constriction.

In support of this Petak et al. [5] demonstrated that during severe constriction due to M c h

that increases in tissue damping, not mirrored by increases in H were related to peripheral

inhomogeneity. Peripheral inhomogeneity is unlikely to be the cause of the tissue responses

demonstrated in these results, as the high lung volume at which the Zrs spectra are

determined would act to split the peripheral airways open and hence reduce the possibility

of peripheral airway inhomogeneity. Using the optimal ventilation waveform technique,

between 0.156-8.1 Hz, Kaczka and coworkers [200] reported that mild constriction,

induced by inhaled M c h caused increases in R a w that accounted for the majority of the

increase in Rrs. These investigators also showed that during bronchoconstriction associated

with a decrease in FEVi of 2 0 % , R a w and G increased but that there was little or no

response in H [200]. The authors fitted Zrs to 2 models of the lung that allowed for, firstly

peripheral airway inhomogeneity and secondly, airway wall compliance to account for

airway wall shunting. Using these models the investigators were able to show that the

majority of the response in Rrs during severe constriction was due to changes in R a w not

G, and that the airway wall shunting occurred causing artifactual increases in G. The O V W

technique used by Kaczka and coworkers [200] allows estimation of the frequency response

of the lung while delivery a waveform sufficient to maintain gas exchange and thus can

136

provide low-frequency Zrs without a pause in breathing. The higher transrespiratory

pressures used in the current study to obtain Zrs may reduce the effects of airway wall

shunting, by the process of splinting the central airways open. So while it is possible that

airway wall shunting may have occurred in current study, causing an artifactual increase in

G, the technique used should act to minimize the influence of airway wall shunting. A

mechanism of disassociated parenchymal response has been suggested by Fredberg and

coworkers [201] w h o showed in isolated lung tissue strips that the time to peak response,

following administration of an agonist, was significantly shorter in rj than in elastance, with

the response in tissue resistance (R) being governed by the product rjE. Using an open-

chest rat preparation Salerno et al [202] demonstrated that the response in Rti could be

attributed to either El or TJ, depending upon the stimulus used. Following inhaled M c h and

changes in lung volume, El wholly accounted for changes in Rti, whereas intravenous M c h

resulted in increases in Rti predominantly attributed to TJ. The early response in r\ has been

attributed to activity in the rapidly cycling cross bridges in the parenchyma, while the

changes in elastance are regulated by the slower cycling cross bridges [201]. This

disassociation between the hysteresivity and elastance of the lung tissues may be responsible

for the changes seen in G within the present study and may be related to increases in r|. It

was not the aim of this study to characterize the time course of the respiratory mechanics

and further studies characterizing this phenomena in infants are required.

The role of deep inhalations (DI) on induced bronchoconstriction has been investigated in

both normal and asthmatic adults [203,204]. Airway hyperresponsiveness was found to be

reduce in normal adults following DI, but was unaffected in mild asthmatic patients [203].

A n effect of DI on M c h responsiveness in infants has yet to be shown, either in normal

[131] or wheezy infants [205]. If a DI effect on the Mch-induced constriction was expected

w e would have expected both FEV0.5 and the oscillatory mechanical parameters to decrease

with successive measurements. This was not the case in the current study, either within all

subjects or within the subgroups of patients. A s the low-frequency F O T requires high lung

volumes to induce a pause in breathing there may be an effect of lung volume on the Mch-

induced constriction.

7.5.4 Conclusions

137

The present study design was similar to that used by Hayden and coworkers [131] to

compare the R T C and R V R T C techniques, however neither of these techniques are affected

by the lack of apnoeic pause required to determine Zrs. A change in methodological design

is required to further assess the usefulness of the low-frequency F O T to detect differences

between patient groups. In conclusion, the present study demonstrated that the low-

frequency F O T is able to detect a response to inhaled methacholine in sedated infants. This

response was attributed predominately to the airways. The provocation concentration

required to cause a 3 0 % increase in R a w is similar to that causing a decrease in FEV0.5 of

1 5 % and would be a suitable outcome variable in further investigations.

138

fi

139

CHAPTER 8: GENERAL DISCUSSION

This thesis has presented the findings of a series of experiments designed to assess the role

of the low-frequency forced oscillation technique (FOT) in measuring infant lung function.

The aims of the thesis were to determine the ability of the low-frequency F O T to non-

invasively partition the respiratory system into lung and chest wall components. This was

initially completed in a rodent model before being translated into a clinical setting. Various

methodological issues were examined in spontaneously breathing infants. Zrs was

partitioned into nasal and lower respiratory impedances. Studies were conducted to establish

which form of forcing signal allowed the optimal balance between maximising signal-to-

noise ratios (S/N) and success rate, while rninimising errors due to cross-talk and harmonic

noise. The relationship between the mechanical properties of the airways and respiratory

tissues and length was determined in a cross-sectional study of normal infants. The airway

and respiratory tissue mechanics in wheezy infants were compared with those of normal

infants and the physiological site of recurrent wheeze in infants determined and lastly the

methacholine responsiveness in infants was assessed using the low-frequency F O T .

8.1 Methodological issues

8.1.1 Measurement of Zrs

The low-frequency F O T has been utilised to obtain Zin from sedated infants during a pause

in breathing. In order to induce this pause in breathing the Hering-Breuer reflex has been

invoked using a computer controlled pump to inflate the infant's lungs three successive

times, after which the airway was occluded at Prs of 20 c m H 2 0 . During the pause in

breathing a forcing signal could be applied to the airway opening and the Zrs obtained (see

Section 2.2.2.1 for full details). The respiratory mechanics presented in this thesis need to

be examined in context of the higher lung volume at which they are obtained. Measurements

at F R C , or perhaps moderate inflation would better represent the conditions governing

spontaneous breathing and facilitate comparisons between the data presented in this thesis

and from other investigators. Petak et al. [76], however demonstrated that in order to

reliably and repeatedly induced a Hering-Breuer reflex, that allowed a pause in breathing of

sufficient length (i.e. > 6 s) occlusion at 20 c m H 2 0 was more practical. The other major

consideration in comparison of the data in this thesis with other studies reporting oscillatory

140

mechanics in infants is the almost distinct frequency ranges used. The studies within this

thesis have uniformly used forcing signals ranging from 0.5-20 Hz, while the majority of

investigators use a frequency range of 4-32 Hz, or in the case of high-frequency oscillations

2-256 Hz. Indeed, current recommendations issued by the European Respiratory Society,

are a frequency range of at least 4-32 H z or if possible 2-48 H z be used [206]. These

frequencies however can only provide information on the mechanical properties of the

airways, and in general investigators using these frequencies report resistance and reactance

values at specific frequency components, or the mean of the components. The differences

in measured frequency ranges make direct comparisons of data obtained in this thesis with

previous investigations difficult. Measurements obtained by superimposing the oscillatory

signal on the spontaneous breathing, at the airway opening, have exhibited a negative

frequency dependence in Rrs [182,197]. This is in direct contrast to studies conducted

within the same frequency range in intubated neonates [126] or in Rrs measurements

determined using the head generator technique [122], indicating a significant influence of

an upper airway shunt in measurements at the airway opening in spontaneously breathing

patients. Numerous studies have been conducted comparing Zrs measured by applying

oscillations directly to the airway opening or using the head generator technique

[112,113,122]. These studies have uniformly shown that in the 6 - 20 H z range the negative

frequency dependence of Rrs is a result of the Zua acting in parallel to Zrs. This frequency

dependence increases in the presence of airway obstruction or ventilatory inhomogeneity

[21,122]. This negative frequency dependence seen in Rrs in the above mentioned studies

is unrelated to the negative frequency dependence exhibited by the respiratory tissues. This

can be clearly seen in the schematic representation of the relative contributions of the

airways and respiratory tissues to Zrs seen in Figure 2.6. The relative contributions of Grs

and Hrs to Zrs is negligible by approximately 5 Hz, whereas model simulations of the effect

of Zua on the measured Zrs clearly shown an influence of the upper airway shunt to

frequencies as high as 20 H z [21,122]. The technique used in this thesis to determine Zrs

should act to minimise the effects of an upper airway shunt. The low-frequency F O T applies

the forcing signal during a pause in breathing and hence the respiratory system is not

undergoing large, rapid changes during the measurement period. In addition the Prs at which

the measurements are made should act to split the central and upper airways open, hence

reducing the possibility of loss of flow and hence the underestimation of Zrs. Studies in

infants using similar frequencies ranges [7,76,123] as those presented in this thesis, have

141

demonstrated the same characteristic low-frequency negative frequency dependence

exhibited in intubated animal studies [4,5,26] and studies in adults [17,96,200].

8.1.2 Application of the constant-phase model to Zin

The constant-phase model ( C P M ) was first used to separate the pulmonary airway and tissue

mechanical properties by Hantos et al. [26]. The investigators showed in open-chested dogs

that the C P M accurately separated the airway and tissue mechanical properties under

baseline and constricted conditions. The validity of this model has been confirmed in both

open-chested and isolated dog lungs [207] and in open-chested rats [5]. Peripheral airway

inhomogeneities resulting from severe constriction, or ventilation mismatching have been

demonstrated to cause increases in G, not related to actual changes in tissue mechanics,

these changes were also shown to be accompanied by a rapid fall in law [5,27,28], similar

artifactual changes have also be shown to result from central airway shunting [27,200]. In

this thesis the C P M has been applied to Zrs, Zl, Z w , and Zlrs and was able to reliably

describe the mechanics properties of these components, additionally the mechanical

characteristics of the nose could be calculated using the derived mechanical properties of Zrs

and Zlrs. The lung impedances measured in Chapter 4 in intubated, mechanically ventilated

infants and young children were found to be accurately described by the C P M , exhibiting

similar frequency dependencies reported in both open- and intact-chest animal studies

[5,149]. The mechanical behaviour of the chest wall tissues was well described by the tissue

compartment of the C P M and contributed approximately a third of the Grs and Hrs

(38.7±7.3% and 34.4±7.4%, respectively). In agreement with studies in cats [4], rats [12]

and human studies [12] the chest wall exhibited a small, but detectable Newtonian resistance

at higher frequencies. In spontaneously breathing sedated infants Zrs was partitioned into

nasal and lower respiratory system components. The C P M could be accurately fitted to Zlrs

and showed similar mechanical properties as seen in the mechanically ventilated patients,

once methodological differences were accounted for (e.g. lung volume and intubation). The

mechanics properties of the nasal pathway were calculated from the derived mechanical

parameters of Zrs and Zlrs (e.g. R n = Rrs - Rlrs). In agreement with previous studies using

posterior rhinometry [151] and forced oscillation above 8 H z [120] R n was found to

constitute a significant proportion of Rrs (-45%). The nasal inertance dominated Irs,

contributing approximately 7 2 % and accounting for the frequency dependent behaviour of

Xrs above 10 Hz. The mechanical behaviour of the respiratory system and it's constitutive

142

components was well described by the parameters of the C P M in all cases.

The input impedance (Zin) in spontaneously breathing infants can be considered as a

combination of series and parallel components as shown in Figure 8.1.

Zin V

Zpnt

Prs

Zn

Zua

Zrs

Patm

Figure 8.1: The input impedance (Zin) of the infant and measurement system during low-frequency oscillation

applied at the nasal opening, via a face mask. The upper airway impedance (Zua) and Zrs are in parallel with each

other and in series with the nasal (Zn) and pneumotachograph (Zpnt) impedances. Flow (V'), oscillatory pressure

(Prs) and atmospheric pressures 0?atm) are also shown.

As previously stated the impedance of the upper airways (Zua) should be minimal in the

measurements of Zin in this thesis due to both the higher lung volumes and the pause in

breathing used in which to obtain Zrs spectra [7,208]. Hence the parallel pathway in Figure

8.1 can be collapsed so that Zin is represented by a series pathway including Zpnt, Zn, and

Zrs. The respiratory system includes the lung and chest wall components in series with each

other and hence can be represented as Figure 8.2.

Zin

Patm

Figure 8.2: Zin represented as a series construction of Zpnt, Zn, Zl and Zw.

If we now include the mechanical behaviour of each component as it has been described by

the CPM, then the low-frequency Zin in infants at a transrespiratory pressure of 20 cmH20

can be portrayed as Figure 8.3.

143

Zin can be represented by a series combination of Zn, Zl and Zw. The frequency independent

resistive properties of the airways are approximately equally split between the nose and the

lung, while the inertive properties of Zin are predominantly determined by the nasal

structures. The majority of the constant-phase description of the respiratory tissues was

attributed to the pulmonary parenchyma. The mechanical behaviour of the respiratory

system, in spontaneously breathing infants can n o w be described in terms of the constant-

phase model as proposed by Hantos et al. [26].

Zin

Rn In Raw law

Gl/

Zn Zl Zw

Figure 8.3: Representation of the partitioning of Zin, into Zn, Zl and Z w compartments. The relative sizes of the

individual mechanical parameters indicates their relative contributions to the overall behaviour of the respiratory

system as a constant-phase model.

8.2 Findings of this thesis in context to the existing literature

Using an esophageal catheter, Zrs was partitioned into lung (Zl) and chest wall (Zw)

impedances, in a Brown Norway (BN) rat model, hence allowing the characterisation of the

airway and parenchymal mechanics and the relative contributions of the chest wall and lung

mechanics to the respiratory system mechanics to be determined [12]. Furthermore, it was

carried out in the context of three longitudinal measurements over a period of 14 days,

allowing animals to act as their own controls. Previously studies investigating the role of the

pulmonary parenchyma have required the use of capsules, glued to the lung surface, to

measure alveolar pressure. More recently the low-frequency F O T has been used in open-

chested animals to obtain the impedance of the lung (Zl), airway and lung tissue mechanics

were subsequently obtained by fitting the C P M to Zl [5,149]. While the use of the low-

frequency F O T has minimised the complexity of the experiment it still requires the thorax

to be opened. Oostveen and co-workers have recently described the use of a two-

compartment whole body plethysmographic technique that allowed the transfer impedance

to be obtained in conscious rats [144], this model was then subsequently adapted for use in

144

mice [145]. The frequency range used in these studies (16 - 208 Hz) does not allow the

mechanical behaviour of the respiratory tissues to be quantified, as was desired in the

present study. The introduction of this methodology into future studies will confer several

advantages. The numbers needed to detect significant changes in lung function will decrease,

as the variability in intact-chest experiments is lower than that reported in open-chest

procedures, studies involving time courses will be able to use a single group of animals

tested at multiple time points, rather than multiple groups tested a single time point. The

development of a technique able to track longitudinal changes in the mechanical properties

of both the airways and lung tissues will allow the long term progression of allergic disease

or novel treatments to be monitored, providing a more realistic model of lung physiology

that will allow a more accurate translation of results from animal studies to human studies.

In a recent development Preuss and co-workers [209] extended the technique presented in

Chapter 3 [12] to allow the measurement of airway and parenchymal mechanics in non-

paralysed rats. .The investigators were able to show that by increasing the ventilation rate

of the animals that reliable estimates of Zl could be obtained and that baseline mechanics

were not altered over a time period of 3 days. In addition the authors demonstrated that the

introduction of both a M c h dose response curve and/or a bronchoalveolar lavage did not

alter baseline mechanics, while the absence of neuromuscular blockade increased the

responsiveness to both inhaled and intravenous M c h [209]. So it can be seen that the

development of a non-invasive technique to monitor the mechanical behaviour of both the

airways and pulmonary tissues has already begun to lead to significant advances in the

development of a more realistic animal model of asthma, or the long-term monitoring of

disease progression or treatment efficacy. This non-invasive methodology was then

translated into a clinical setting in 5 mechanically ventilated patients undergoing cardiac

surgery (Chapter 4). The low-frequency F O T was able to reliably partition the Zrs into Zl

and Z w components. The C P M could then be fitted to these impedance spectra and the

mechanical properties of the airways and tissues described. The chest wall was found to

contribute a negligible amount to the respiratory system resistance (Rrs) and inertance (Irs),

but to contribute a significant amount to the respiratory tissue damping (Grs) and elastance

(Hrs) (38.5±7.3% and 34.4±7.4%; mean±SEM, respectively). Previous studies attempting

to partition respiratory mechanics in infants or children have reported values of lung

resistance and elastance using passive or dynamic techniques, or values of airway resistance

using the interrupter technique. However these techniques are unable to provide

145

simultaneous assessments of both the airway and lung tissue mechanics. In addition pre

existing techniques are frequency dependent making comparisons between age and patient

groups difficult. The introduction of a non-invasive technique that allows the frequency

dependent characteristics of the lung tissue to be determined will allow more detailed

investigations into the development of parenchymal disease in infants and young children.

The initial study, characterising the lung and chest wall mechanics in intubated ventilated

patients was modified and extended to partition Zrs into Zn and Zlrs impedances in sedated,

spontaneously breathing infants (Chapter 4). Z n was found to influence only the airway

properties of the respiratory system, contributing 44.6±4.9% (mean ± S E M ) of the Rrs and

a majority of the Irs (71.7±3.5%). These relative contributions were also found to be

constant with increasing length, indicating that changes in the nasal mechanics track those

of the respiratory system, during the first two years of life. The findings of these studies have

allowed the relative contributions of the anatomical structures of the respiratory system to

the low-frequency respiratory system mechanics to be described for the first time. These

findings and their contributing technical factors have been outlined previously in detail in

Section 8.1.2. Future infant lung function studies using a nasal or esophageal catheter will

be able to improve the sensitivity of the technique depending on whether the airways (nasal

catheter) or pulmonary tissues (esophageal catheter) are the primary site of interest. This

will allow for the first time an accurate characterisation of the development of the low-

frequency mechanics of the lung chest wall and respiratory system in sedated infants. The

use of such techniques in infants with a history of wheeze or chronic lung disease will allow

information on the role of the airways and pulmonary parenchyma on the development of

such diseases and the possible impact of treatments within such patient groups, to be

obtained.

The work presented in this thesis has reported the mean coefficients of variation of the

mechanical parameters of the airways (Raw and law) and respiratory tissues (Grs and Hrs)

(Chapter 5). The intra-subject variability of the mechanical parameters of the airway and

respiratory tissues compared well with values reported previously in the literature (Table

5.3). Forced expiratory volumes (FEV0.5) and low-frequency F O T mechanical parameters

were determined in a cross-sectional study of infants with no history of respiratory disease.

The relationship between the lung function parameters and length was investigated using

multivariate regression analysis. The effects of gender, passive smoke exposure and family

146

history of asthma on the lung function parameters were also investigated. Both airway and

tissue parameters demonstrated a decreasing quadratic relationship with increasing length,

while FEV0.5 showed an increasing cubic relationship with length. A family history of asthma

was found to have a negative effect on Raw, H and FEV0.5, whereas gender did not have any

effect. The numbers of infants exposed to environmental tobacco smoke were too small to

analyse accurately. The establishment of normal ranges are essential if comparisons are to

be made in patient groups with a history respiratory disease. Previous studies using the

forced oscillation technique to report normal values of respiratory mechanics have used a

wide choice of frequency ranges and numerous different techniques and are not suitable for

comparison with results generated using the low-frequency F O T . While some initial studies

have been carried out using the low-frequency F O T [7,76,123] issues such as repeatability

and intra-subject variability have not been addressed. The initial work presented in Chapter

5 of this thesis will allow detailed studies into the physiological processes associated with

the development of respiratory disease to be carried out in the future. Longitudinal studies

are needed if the role of gender, family history of asthma and passive smoking on airway and

tissue mechanics are to be accurately defined. Previous studies have documented effects of

these factors on lung function, but the results are not clear and in many cases inconsistent.

The rate of change in airway mechanics with length was slower than that of the tissue

mechanics (both G and H ) supporting the concept of dysanaptic growth within this

population. Previous studies have both confirmed [61,176] and refuted [177] that the

concept of dysanaptic growth applies to the growth of the respiratory system. Hibbert and

co-workers [177] proposed that as lung volumes, maximal flows and pressures all track with

growth then dysanaptic growth of the airways and respiratory tissues cannot occur. These

investigators however studied an older population (8-12 years of age), by which time it is

certainly feasible that the majority of growth had occurred and that further growth would

be proportional. These previous studies have been hampered by the fact that the mechanical

behaviour of the airways and respiratory tissues have not been assessed directly. Lanteri et

al. [61] did examine the airway and tissue mechanics, however the population studied was

in intubated infants and children undergoing surgery and does not accurately represent a

healthy population. The low-frequency forced oscillation technique offers an opportunity

to characterise the growth of both the airways and the respiratory tissues in a healthy infant

population. Furthermore the technique has the possibility to identify if infants with

147

respiratory disease, such as chronic lung disease or cystic fibrosis, have differing growth

patterns and the site at which such differences may occur.

A further cross sectional study was carried out in 22 asymptomatic infants with a history of

recurrent or persistent wheeze (Chapter 6). Baseline airway and respiratory tissue mechanics

were determined and standardised variants (Z scores) were used to compare the wheezy

population to the previously described normal population, hence ascertaining the role of the

airways and respiratory tissues in the development of chronic wheeze in infancy. The

mechanical properties of the airways and respiratory tissues were found to be significantly

abnormal in wheezy infants older than 1 year of age, when compared to their normal

counterparts. Suggesting that infantile wheeze is not merely related to airway function, but

rather an alteration in mechanical characteristics of both the airways and the respiratory

tissues. Previous studies have not consistently shown differences in lung function parameters

in infants with a history of recurrent wheeze. Studies have shown altered Rrs and Crs [181]

and V m a x F R C [178] in wheezy compared to healthy infants. In contradiction to these

studies tidal volume, Crs and respiratory system conductance [178], F R C [181] and Raw,

specific R a w and thoracic gas volume ( T G V ) [183] have all been shown to be unaltered in

infantile wheeze. The present study demonstrated abnormal respiratory system mechanics

in wheezy infants older than one year indicating that ongoing recurrent wheeze is driven by

underlying alterations in airway and lung tissue physiology. It is possible that the

introduction of an esophageal catheter to allow direct measurement of the lung would show

abnormalities at an earlier age than demonstrated in this thesis. These results reinforce the

importance of determining the mechanical behavior of the respiratory tissues in infant lung

function studies.

The responsiveness to inhaled methacholine (Mch) in 17 infants was assessed using the low-

frequency F O T , by comparing the threshold concentration in R a w with a 1 5 % drop in

FEV0.5 (PC15FEV0.5). Doubling concentrations of inhaled M c h were administered until the

maximal dose (16 mg/mL) was achieved or a response in FEV0.5 recorded. A n increase in

R a w of at least twice the baseline variability (2 times coefficient of variation; T C R a w ) was

considered to be an positive response as determined by the low-frequency F O T . The T C R a w

was found to be significantly lower than that of the PC15FEV0.5. Upon further examination

this was attributed to the relative changes required to reach the respective responses. Using

148

previously determined baseline variability in FEV0.5 of approximately 5-7% [47,131] a

change in FEV0.5 of 1 5 % represents an increase between 2 and 3 times that of baseline

variability, while the change required to achieve T C R a w is by definition, twice that of the

baseline variability. A n increase in R a w of 3 0 % equates to an increase of approximately 2.7

times that of the baseline variability (Chapter 5) and would be equivalent to PC15FEV0.5.

Indeed, PC3oRaw and PC15FEV0.5 were found to be the same (1.5; 0.5-2 m g / m L and 1.5;

1-4 mg/mL (median;25-75% CI), respectively), indicating that the low-frequency F O T is

effective in detecting a response to inhaled M c h in infants. Significant changes in law, G and

rt were also recorded, while no response was noted in H. While central airway shunting may

have contributed to the increase in G and rj, but not H [200], the use of a transrespiratory

pressure of 20 c m H 2 0 should act to splint the central airways open and hence reduce airway

shunting. The disassociation in responses between r| and H has been previously

demonstrated by Fredberg and co-workers [201] and may explain the tissue response to

inhaled M c h seen in the present study. Changes in central airway mechanics were thought

to be responsible for the increase in law, indicating inhaled M c h acts on both the central and

peripheral airways in sedated infants. A number of limitations were identified. Zrs spectra

could not be obtained in some infants undergoing a response to M c h (i.e. FEV0.5 decreased

by 15%) due to increased respiratory rates, wheeze or cough and hence pauses in breathing

of sufficient lengths could not be induced. The chosen challenge protocol exacerbated this

problem. Cumulative M c h inhalations were used with successive doses administered every

5 minutes, during which time both FEV0.5 and Zrs spectra were obtained. In infants with a

strong reaction to a particular M c h concentration, difficulties were experienced in obtaining

both lung function measures. A revised protocol using PC3oRaw as the outcome variable

would allow more time at each inhalation for successful Zrs spectra to be obtained. The

possible disassociation between r\ and H needs to be investigated as does the role of

hyperresponsiveness in infants with a history of wheeze. The use of M c h challenge testing

in infants with a history of recurrent wheeze may provide further information on the role of

the pulmonary tissues in the development of early asthma.

8.3 Future directions

Further studies are needed to determine the full potential of the low-frequency forced

oscillation technique in infant lung function. Longitudinal studies with large numbers are

149

required to examine, in more detail the effect of environmental smoke exposure, gender and

family history on the subsequent development of the airways and respiratory tissues, as well

as which risk factors may be most important in predicting the differences between those

infants with transient and persistent wheeze. The active control of end-expiratory lung

volume in infants may be a major factor in the lack of differences between

hyperresponsiveness in infants with and without wheeze [178]. The combination of the low-

frequency F O T and measurements of lung volume may provide more detailed information

on the mechanical behaviour of both the airways and respiratory tissues in infants, allowing

better discrimination between healthy and diseased infants. The additional inclusion of

catheters to partition Zrs will also increase the sensitivity of the technique to detect changes

in R a w by 1 0 0 % and in Gl and HI by 5 0 % , depending upon the location of the catheter

(nasal and esophageal, respectively). Based on the data presented in this thesis either a nasal

or esophageal catheter should be routinely used, depending upon whether the airways or

respiratory tissues are the primary site of interest. Future investigations into the

responsiveness of infants to inhaled agonists using the low-frequency F O T should use

PC3oRaw as the outcome variable. The redesign of the challenge protocol will increase the

successfulness of the technique and allow more detailed information regarding

hyperresponsiveness in infants. Further information on the mechanical changes in the

pulmonary parenchyma during bronchoconstriction should be obtained, with studies

investigating the disassociation between the mechanical responses in G, H and r| to inhaled

M c h warranted. Additional studies in symptomatic infants are needed, allowing comparisons

with both healthy and asymptomatic patient groups. Such studies may provide valuable

information on the mechanical changes in either the airways or pulmonary parenchymal

accompanying episodes of viral induced wheeze, exacerbations of chronic wheeze or cough

as well as other respiratory problems, such as chronic lung disease.

8.4 Conclusions

This thesis has detailed much of the initial work required to allow the low-frequency forced

oscillation technique to be adopted for investigations into the genesis of respiratory disease

in infants. The anatomical components of the respiratory system were characterised. The

nose was found to contribute in the order of half of the Rrs and a majority of the Irs, the

chest wall significantly influenced the constant-phase tissue mechanics of the respiratory

system contributing approximately a third of the measured Grs and Hrs. Regression

150

equations establishing the relationships between both airway and respiratory tissues

mechanics and length were established. Family history of asthma was found to have a

significant detrimental effect on Raw, H and FEV0.5. These regression equations were then

used to demonstrate that infants older than 12 months of age with ongoing recurrent wheeze

have abnormal airway and respiratory tissue mechanics. The ability of the low-frequency

F O T to detect hyperresponsiveness in infants to inhaled M c h was shown with significant

responses seen in Raw, law, G and r\, but not in H. The adoption of PC 3oRaw as a suitable

outcome variable was recommended for use in future studies.

In conclusion, the observations of Sly and colleagues [7] on the low-frequency forced

oscillation technique in infants have been addressed. The technical factors associated with

use of the technique in future studies have been evaluated, allowing investigations into

clinical issues in patient groups to be carried out.

151

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