ERS International Congress Amsterdam

26–30 September 2015

Postgraduate Course 7 Basic respiratory mechanics

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Saturday, 26 September 2015 09:30–13:00

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Postgraduate Course 7 Basic respiratory mechanics

AIMS: The mechanics of breathing is best described in terms of airways resistance and lung compliance. However, these measurements are not readily available in clinical practice and instead clinicians must focus on the indirect information available from spirometry. The great advantage of spirometry is its known immediate reproducibility and limited between-test variation in both healthy and diseased settings. These properties, coupled with its extensive use in clinical decision making, ensure that spirometry will be used to manage respiratory disease for the foreseeable future. One particular advantage of spirometry is the availability of age- and height-related normal values against which any recording can be judged. This PG course on basic respiratory mechanics will help clinicians to better understand and interpret lung function test results. TARGET AUDIENCE: Pulmonologists, respiratory therapists, respiratory physicians, clinical researchers, general practitioners, research fellows, nurses, and trainees.

CHAIRS: M. Polkey (London, United Kingdom), S. Verges (Grenoble, France) COURSE PROGRAMME PAGE

09:30 Modelling respiratory mechanics: a multi-scale approach 5 J. Bates (Burlington, United States of America)

10:15 Measuring respiratory mechanics: invasive and noninvasive systems 53 A. Aliverti (Milan, Italy)

11:00 Break

11:30 Using respiratory mechanics to explain and interpret spirometry 117 J. Gibson (Newcastle upon Tyne, United Kingdom)

12:15 Using respiratory mechanics to study exercise limitation in disease 164 L. Romer (Uxbridge, United Kingdom)

Additional course resources 200

Faculty disclosures 201

Faculty contact information 202

Answers to evaluation questions 203

Th e ERS Handbook of Respiratory Medicine is a concise, compact and easy-to-read guide to each of the key areas in respiratory medicine. Its 18 chapters, written by clinicians and researchers at the forefront of the fi eld, explain the structure and function of the respiratory system, its disorders and how to treat them.

Now in its second edition, the Handbook is a must-have for anyone who intends to remain up to date in the fi eld, and to have within arm’s reach a reference that covers everything from the basics to the latest developments in respiratory medicine.

Accredited by EBAP for 18 hours of European CME credit

To buy printed copies, visit the ERS Bookshop at the ERS International Congress 2015 (Hall 1, Stand 1.D_12).

Electronic: WWW.ERSPUBLICATIONS.COMPrint: WWW.ERSBOOKSHOP.COM

EDITED BY PAOLO PALANGE AND ANITA K. SIMONDS ISBN 978-1-84984-040-8

THE ERS HANDBOOK OFrespiratory medicine

Lung Mechanics: Multi-scale Modeling

Dr Jason H.T. Bates University of Vermont

College of Medicine Burlington Vermont 05405-007

UNITED STATES OF AMERICA jason.h.bates@uvm.edu

AIMS

Demonstrate how simple inverse models are used to interpret measurements of lung impedance. Demonstrate how more complicated forward models can be used to test hypotheses about

physiologic mechanism. Show how the forward-inverse modeling paradigm can be used to understand airway

hyperresponsiveness. SUMMARY Assessing Lung Mechanics using Inverse Models Performing a medical diagnosis of any kind requires probing the human body in some manner in order to find out something about its inner workings. In fact, this is a paradigm that is ubiquitous in science; the system under investigation is disturbed in some way by being subjected to a controlled perturbation known as an input. The investigator then observes the system’s reaction, known as the output, and tries to draw conclusions about mechanisms internal to the system that are not directly observable. This paradigm applies to the two standard approaches to clinical assessment of lung function, spirometry and plethysmography, and in both cases the inputs are generated by the subjects themselves. In the case of spirometry the inputs consist of forced expiratory maneuvers that produce the outputs FVC and FEV1. In the case of spirometry the inputs are typically rapid shallow panting maneuvers that produce outputs of pressure and flow measured at various sites within the plethysmograph chamber. Both spirometry and body plethysmography have relatively recent histories in the annals of medicine, coming into common usage during the middle of the 20th Century when the sensors and recording devices needed to gather the data they generate became available. This same technology allowed the advent of a third approach to pulmonary diagnosis, an approach known as the forced oscillation technique (FOT). The FOT has a fundamental difference to both spirometry and body plethysmography in that it does not require the subject to be involved in the measurement process in an active way. Instead, perturbations are applied to the lungs by an external device that generates oscillations in air flow which are directed into the airway opening (e.g., the nose and/or mouth in an ambulatory subject or the tracheal opening in a mechanically ventilated patient). It is intuitively obvious that the pressures generated at the airway opening as a result of applied oscillations in flow must contain at least some information about lung function. For example, the increased airway resistance typical of asthma would be expected to give rise to an increased swing in pressure for a given amplitude of oscillating flow. Similarly, a subject with pulmonary fibrosis would be expected to require an elevated swing in pressure when receiving a given oscillating volume. Putting this kind of information in quantitative form and interpreting it in physiological terms, however, requires mathematical analysis derived from the fields of engineering and physics. This is unfamiliar territory to many in the biological and medical realms, which may explain why the FOT is only recently beginning to gain acceptance as a technique for clinical assessment of lung

5

function even though its origins are as old as its established cousins, spirometry and body plethysmography. On the other hand, FOT has some major advantages over these other two techniques in terms of its ability to link structure to function in the lung. Resistance, Elastance and Inertance The most basic model of lung mechanics, the one that is most commonly invoked by physiologists, represents the lungs as a single airway conduit connected to a single elastic alveolar unit (Fig. 2.1). There are clear advantages to such a simple description. For example, this model makes it possible to talk about airway resistance in an unambiguous manner because a conduit of given length and circular cross-section has a well-defined airflow resistance (R) given by the ratio of the pressure difference (P)

between its two ends to the flow (V ) through it thus:

V

PR

(1)

R is a reflection of how difficult it is to drive air through the conduit physiological and is therefore a measure of physiological function. Physical theory allows R to be related to the dimensions of the conduit and the properties of the gas. In other words, one can use a measurement of R to infer something about the dimensions of the pulmonary airways to the extent that the conducting airway tree can be usefully represented as a single conduit. Similarly, the recoil characteristics of the lungs are embodied in the single parameter E, a measure of elastic stiffness. Relating the magnitude of E to underlying mechanism is not as straightforward as described above for R because the physical processes involved are more complicated and less well understood. When the lung is inflated elastic energy is stored in the protein fibers that comprise the parenchymal tissue as well as in the air-liquid interface as a result of surface tension, and their relative contributions change with volume and volume history. Nevertheless, the value of E tends to change in predictable ways with certain lung diseases. For example, when pulmonary fibrosis leads to the appearance of aberrant connective tissue in the parenchyma the value of E increases, whereas when the parenchyma becomes destroyed in emphysema E decreases. If flow is oscillated into the lungs at high enough frequencies then a significant fraction of the pressure driving the oscillations is needed to accelerate and decelerate the gas in the airways. The mass of this gas thus has an inertance, I. The complete equation of motion of the model shown in Fig. 1 is thus

)()()()( tVItVRtEVtP (3)

E

R

Pel

Pao

V

V

Figure 1: A single alveolar compartment is represented as a pair of telescoping canisters, and is connected to the environment by a single conduit representing the conducting airway tree. A spring connected between the two canisters generates an elastic pressure, Pel, inside the alveolar compartment as its volume, V, increases due to the

accumulation of flow, V . The conduit has a flow resistance R given by Eq. 2, where P is the difference between the pressure at the airway opening (Pao) and Pel.

6

Lung Impedance The mechanical input impedance of the lung is a general measure of the difficulty with which air can be forced into it. It is expressed mathematically as the ratio of the Fourier transform of P(t) to the Fourier

transform of )(tV (both computed using the fast Fourier transform algorithm) as follows:

)()(

)(fV

fPfP

(4)

where the argument of frequency (f) denotes the Fourier transform of the corresponding time (t) domain signal. The impedance of the model shown in Fig. 1, determined by the taking the Fourier transform of Eq. 3, is

f

EfIiRtZ

22)( (5)

This predicts that the real part of impedance (known as resistance) is equal to a constant, R, and thus

does not change with f. The imaginary part of impedance, which that part multiplied by 1i , is known as reactance and this does change with f. Importantly, reactance is zero at the resonant frequency, fres, when 2fI = E/2f, giving

I

Ef res 24

(6)

In reality, R and E also both depend on f because the lung tissue is viscoelastic. Consequently, a model that typically provides much better fits to experimental measurements of Z(f) between 1 and 20 Hz is the so-called constant-phase model with an impedance given by [1]

)2(2)(

f

iHGfIiRtZ N (7)

RN is a Newtonian resistance that has been shown to closely approximate the flow resistance of the conducting airways [2], and I is inertance. G is called tissue damping and is a measure of the rate at which energy is dissipated in the form of heat when the lung tissues are oscillated, and H is a measure of tissue stiffness closely related to E. In other words, the spring in Fig. 1 is replaced with a constant-phase tissue unit characterized by the parameters G and H. The parameter α is determined by the values of G and H [1], so the impedance of the constant-phase model has only 4 adjustable parameters – RN, I, G and H – and in mice one can neglect the effects of I below 20 Hz. Measurements of Z(f) in animal models are often used to evaluate RN, G and H which are then interpreted as follows. An increase in RN typically means that airway caliber has decreased. When G and H increase in the same relative proportions it indicates that a fraction of the lung has become decrecruited. If G increases proportionately more than H it is suggestive of the development of regional heterogeneities in lung function.

7

Simulation Models of Lung Mechanics The models described above that are used for assessing lung mechanical function from experimental measurements of pressure, flow and volume are called inverse models and are necessarily simple because a model with too many parameters cannot be uniquely identified from a given data set; the numerous parameters of complicated models cannot be evaluated within closely defined limits and there are usually a variety of ways in which the components of complex models can be arranged that all fit the data equally well. However, models can also be used to simulate pressure, flow and volume data based on detailed a priori knowledge of lung structure. Such models are called forward models and are useful for testing hypothesis about possible mechanisms underlying the experimental findings embodied in the parameters of inverse models. Forward models have been used, for example, to infer underlying mechanisms of airways hyperresponsiveness in animal models of asthma [3-6] and to understand the dynamics of recruitment and derecruitment in animal models of acute lung injury [7, 8]. The use of forward models to test a hypothesis can never lead to definitive acceptance of the hypothesis because one never knows for sure if the model contains all the necessary detail to allow such a conclusion. However, forward models can be very useful for helping to establish the feasibility of a particular hypothesis, and in particular may allow a hypothesis to be refuted. REFERENCES 1. Hantos, Z., et al., Input impedance and peripheral inhomogeneity of dog lungs. J Appl Physiol

(1985), 1992. 72(1): p. 168-78. 2. Tomioka, S., J.H. Bates, and C.G. Irvin, Airway and tissue mechanics in a murine model of asthma:

alveolar capsule vs. forced oscillations. J Appl Physiol (1985), 2002. 93(1): p. 263-70. 3. Bates, J.H., et al., The synergistic interactions of allergic lung inflammation and intratracheal

cationic protein. Am J Respir Crit Care Med, 2008. 177(3): p. 261-8. 4. Bates, J.H., et al., Exaggerated airway narrowing in mice treated with intratracheal cationic protein. J

Appl Physiol (1985), 2006. 100(2): p. 500-6. 5. Wagers, S., et al., The allergic mouse model of asthma: normal smooth muscle in an abnormal lung?

J Appl Physiol (1985), 2004. 96(6): p. 2019-27. 6. Wagers, S.S., et al., Intrinsic and antigen-induced airway hyperresponsiveness are the result of

diverse physiological mechanisms. J Appl Physiol (1985), 2007. 102(1): p. 221-30. 7. Allen, G. and J.H. Bates, Dynamic mechanical consequences of deep inflation in mice depend on

type and degree of lung injury. J Appl Physiol (1985), 2004. 96(1): p. 293-300. 8. Massa, C.B., G.B. Allen, and J.H. Bates, Modeling the dynamics of recruitment and derecruitment in

mice with acute lung injury. J Appl Physiol (1985), 2008. 105(6): p. 1813-21. 9. Bates, J.H., et al., Oscillation mechanics of the respiratory system. Compr Physiol, 2011. 1(3): p.

1233-72. 10. Bates, J.H.T., Lung mechanics: An inverse modeling approach. 2009, Cambridge: Cambridge

University Press. 11. Lauzon, A.M., et al., A multi-scale approach to airway hyperresponsiveness: from molecule to organ.

Front Physiol, 2012. 3: p. 191.

8

EVALUATION 1. Inverse models of lung mechanics necessarily must simple (i.e. have few adjustable parameters)

because a. A given data set typically will not allow many parameters to be evaluated within narrow

limits. b. Having too many different parameters can make it difficult to understand their various

physiological interpretations. c. There are invariably many different plausible ways that the numerous compartments of a

complicated model can be arranged, and there is no way of knowing from the data which one is the best.

d. All of the above.

2. Resonant frequency occurs when the imaginary part of impedance (the reactance) is a. Unity b. Zero c. Infinity d. Negative

3. The following are potential mechanisms underlying airway hyperresponsiveness a. Increased mass of airway smooth muscle b. Thickening of the airway epithelium c. Increased access of an inhaled agonist to the airway smooth muscle d. All of the above

4. Complicated forward models of the lung (i.e. simulation models) are most useful for a. proving hypothesis b. fitting to experimental data c. disproving hypotheses d. none of the above

9

MODELING RESPIRATORY MECHANICS: A MULTI-SCALE APPROACH

Jason H.T. Bates, PhD, DSc

Professor of Medicine, Pulmonary/Critical Care Division

University of Vermont College of Medicine

10

Conflict of interest disclosure I have no, real or perceived, direct or indirect conflicts of interest that relate to

this presentation. I have the following, real or perceived direct or indirect conflicts of interest that

relate to this presentation: Affiliation / financial interest Nature of conflict / commercial company name

Tobacco-industry and tobacco corporate affiliate relatedconflict of interest

Grants/research support (to myself, my institution or department):

Honoraria or consultation fees:

Participation in a company sponsored bureau:

Stock shareholder:

Spouse/partner:

Other support or other potential conflict of interest:

This event is accredited for CME credits by EBAP and speakers are required to disclose their potential conflict of interest going back 3 years prior to this presentation. The intent of this disclosure is not to prevent a speaker with a conflict of interest (any significant financial relationship a speaker has with manufacturers or providers of any commercial products or services relevant to the talk) from making a presentation, but rather to provide listeners with information on which they can make their own judgment. It remains for audience members to determine whether the speaker’s interests or relationships may influence the presentation.Drug or device advertisement is strictly forbidden.

X

11

Introduction

AIMS

• Aim 1: Demonstrate how simple inverse models are used to interpret measurements of lung impedance

• Aim 2: Demonstrate how more complicated forward models are used to test hypotheses about physiologic mechanism

• Aim 3: Show how the forward-inverse modeling paradigm can be used to understand airways hyperresponsiveness

12

LinkMathematical or

computational model

Human Lung disease

Mouse Models of Lung disease

FunctionLung Mechanics

StructureLung anatomy

13

ASSESSMENT OF LUNG MECHANICS

14

The Linear Single-Compartment Model of the Lung

VREVPPP el

15

Fit of single-compartment linear model to data from ventilated patient using multiple linear regression

16

Dose-response to methacholine (mice)

Tomioka et al. J Appl Physiol 93: 263-270, 2002

17

)()()( fVfZfP

Lung impedance

Lung(linear dynamic system)

Output

P(t)

Input

)(tV.

)(

)()(

fV

fPfZ

18

Measuring respiratory system impedance in spontaneously breathing subjects

19

Forced oscillation technique in mice using the Flexivent…

20

)()()()( tVItVRtEVtP VP ,

R I (due to mass of gas in proximal airways)

EIiRV

VIiVRVEi

P

)(

)()()()(

F.F.T.

21

E

IiRVP )()(

I

Eres

}

The reactance term is zero when:

22

Lung with Emphysema

old normal lung

normal lung

Verbeken et al., J Appl Physiol72: 2343-53, 1992.

Resonant frequency I

Eres

23

Respiratory impedance in obese asthmatics

0 5 10 15 20 25 30 35

4

5

6

7

8

Control baseline Control post-surgery Asthma baseline Asthma post-surgery

Re

sist

an

ce (

cmH

2O

.s.L

-1)

A

0 5 10 15 20 25 30 35

-3

-2

-1

0

1

2 B

Re

act

an

ce (

cmH

2O

.s.L

-1)

Frequency (Hz)

Al-Alwan et al. Am J Respir Crit Care Med189: 1494-1502, 2014

24

Rc Ic Rp

Ec Ep

Ep

Ec

Rp

RcIc

Interpreting impedance requires a mathematical model

25

0 5 10 15 20 25 30 35

4

5

6

7

8

Control baseline Control post-surgery Asthma baseline Asthma post-surgery

Res

ista

nce

(cm

H2O

.s.L

-1)

A

0 5 10 15 20 25 30 35

-3

-2

-1

0

1

2 B

Rea

ctan

ce (

cmH

2O.s

.L-1)

Frequency (Hz)

Experimental data

0 5 10 15 20 25 30 35

-3

-2

-1

0

1

2 B

Re

act

an

ce (

cmH

2O

.s.L

-1)

Frequency (Hz)

0 5 10 15 20 25 30 35

4

5

6

7

8A Control baseline

Control post-surgery Asthma baseline Asthma post-surgery

Re

sist

an

ce (

cmH

2O

.s.L

-1)

Model fits

26

0

20

40

Eto

tal (

cmH

2O.L

-1)

0

50

100

Ep

(cm

H2O

.L-1)

0

4

8

R

p (c

mH

2O.s

.L-1)

0

100

200

E

c (c

mH

2O.L

-1)

0.000

0.005

0.010

Ic

(cm

H2O

.s.L

-2)

0

3

6

Rc

(cm

H2O

.s.L

-1)

Ep

Ec

Rp

RcIc

27

Constant-phase tissue model

GH

arctanπ

2α

G: tissue resistance

H: tissue stiffness

RN: airway resistance

αNf2

iH-GRZ(f)

Impedance of the constant phase model

28

αNf2

iH-GRZ(f)

RN airway resistance

G tissue damping

H tissue stiffness

RN narrowed airways

G, H in same proportion loss of peripheral lung units

G more than H regional heterogeneities

Interpreting the constant-phase model

29

MODELING AIRWAYS RESPONSIVENESS TO A SMOOTH MUSCLE AGONIST

IN AN ANIMAL MODEL

30

A mouse model of “asthma”

Sensitization:Ovalbumin+Alum

Challenge:Ovalbumin+Alum

31

0 30 60 90 120 150 1800

50

100

150

200

250

Post-DI

Control Inflamed

RN

(% c

hang

e fro

m b

asel

ine)

Wagers et al. J Appl Physiol96: 2019-2027, 2004 32

0 30 60 90 120 150 1800

50

100

150

200

250

Post-DI

Control Inflamed

RN

(% c

hang

e fro

m b

asel

ine)

0 30 60 90 120 150 1800

50

100

150

200

Gt (

% c

hang

e fro

m b

asel

ine)

Post-DI

Wagers et al. J Appl Physiol96: 2019-2027, 2004 33

0 30 60 90 120 150 1800

50

100

150

200

250

Post-DI

Control Inflamed

RN

(% c

hang

e fro

m b

asel

ine)

0 30 60 90 120 150 1800

50

100

150

200

Gt (

% c

hang

e fro

m b

asel

ine)

Post-DI

0 30 60 90 120 150 1800

30

60

90

120

150

Ht (

% c

han

ge fr

om b

asel

ine)

Time (s)

Post-DI

Wagers et al. J Appl Physiol96: 2019-2027, 2004 34

Computational model of mouse lung

18

17 11

1051612

4

8

r

lR

f

iHG

2

35

0 5 10 15 20

-4

-3

-2

-1

0

Frequency (Hz)

Rea

ctan

ce (

cmH

2O.s

.ml-1

)

0.0

0.2

0.4

0.6

0.8

1.0

1.2R=0.29 cmH

2O.s.ml-1

I = 0.0030 cmH2O.s2.ml-2

G = 3.0 cmH2O.s.ml-1

H = 20.4 cmH2O.s.ml-1

Data (+/- SD) Fit

Res

ista

nce

(cm

H2O

.s.m

l-1)

Fit of model to baseline mouse impedance data

36

Control mice:

Airway narrowing alone

0 50 100 150

0.0

0.2

0.4

Time (s)

0.0

0.5

1.0

1.50.0

0.5

1.0

1.5 experimental simulation

Hrs

Grs

Rrs

37

Airway closure by liquid bridging

Lundblad et al. Am J Respir Crit Care Med 171: 1363-1370, 2007

38

Control mice:

Airway narrowing

Airway closure at r = 35m

G/H increasing with airway narrowing

0 50 100 150

0.0

0.2

0.4

0.0

0.5

1.0

1.50.0

0.5

1.0

1.5 experimental simulation

Time (s)

Hrs

Grs

Rrs

39

0 50 100 150

0.0

0.5

1.0

1.5

2.0

Time (s)

H

0.0

0.5

1.0

1.5

2.0

G

0.0

0.5

1.0

1.5

2.0

2.5 Experimental data Simulated data

R

Inflamed mice:

Same model as controls, with increased airway narrowing

40

Control

Inflamed

41

0 50 100 150

0

1

20

1

2

0

1

2 experimental simulation

Hrs

Grs

Rrs

Time (s)

Inflamed mice:

Same degree of airway narrowing as controls

Airway lining increased by 20 m

Airway closure radius increased to 45 m

42

Acutely allergically inflamed mice (BALB/c) are hyperresponsive to methacholine aerosol because of a physical thickening of the airway epithelium and an increase in airway secretions. This results in enhanced closure of small airways, NOT increased shortening of the airway smooth muscle.

43

0 5 10 15 20-4

-2

0

2

4

6

Time (s)

Pre

ssur

e (c

mH

2O)

-0.05

0.00

0.05V

olum

e (m

l)

0

2)(iHG

fIiRfZ

0 5 10 150

1

2

R (

cmH

2O.s

.ml-1

)

Time (s)

44

Cojocaru et al. J Appl Physiol 104: 1601-1610, 2008

Acute Controls

0 5 10 15 200

1

2

PEEP 1 PEEP 3 PEEP 6

R (

cmH

2O.s

.ml-1

)

Time (s)

Control Acute Ova

0 5 10 15 200

1

2

R (

cmH

2O.s

.ml-1

)

Time (s)

Allergically inflamed

0 5 10 15 200

1

2

R (

cmH

2O.s

.ml-1

)

Time (s)

Acute ControlsHigh dose

Controlhigh-dose MCh

45

Acutely allergically inflamed BALB/c mice are hyper-responsive to injected methacholine.

• response in the central airways, not lung periphery

• like normal mice injected with 3 times Mch dose

Acutely allergically inflamed BALB/c mice are hyper-responsive to aerosolized methacholine

• response in both central airways and lung periphery

• due to epithelial thickening and increased secretions

The airway smooth muscle is hyper-responsive

The airway smooth muscle is normal

?

46

airway wall compartment

tissue compartment

airway smooth muscle

compartment

aerosol injection

Epithelial barrier Endothelial barrier

(Bates et al. J Appl Physiol 112: 1670-1677, 2012.)

47

Relaxed Contracted

Reduced airway-parenchymal tethering• Inflammation around airways• Reduced lung volume

Stronger smooth muscle contraction• Increased muscle mass• Increased muscle stimulation

Thicker airway wall • Inflammation • Edema

48

Lauzon et al. Frontiers in Computational Physiology and Medicine3: article 191, 2012.

49

SUMMARY

• Inverse modeling (system identification)– Simple model structure

– Parameter values represent physiological quantities of interest

• Forward modeling (simulation)– Can be very complicated

– Serve as “virtual laboratories” for testing feasibility of specific hypotheses

• Linking structure to function– Human patients with asthma

– Mouse models of “asthma”

50

Conclusions

• Airways responsiveness can be due to a variety of unrelated factors.

• Discerning which factors are at play is often very difficult.

• Computational modeling plays an essential role in helping to properly interpret the physiological basis for observed effects on lung function

51

Acknowledgements:

Minara Aliyeva, MD

Ana Cojocaru, MD

Kara Grant, BS

Charles Irvin, PhD

Anne-Marie Lauzon, PhD

Lennart Lundblad, PhD

Michael Sanderson, PhD

Anne-Marie Lauzon, PhD

Merryn Tawhai, PhD

James Sneyd, PhD

Scott Wagers, MD

Supported by:

HL-62746

HL-67273

HL-87788

RR-15557

52

Measuring respiratory mechanics: invasive and noninvasive systems

Prof. Andrea Aliverti Politecnico di Milano

Dipartimento di Elettronica, Informazione e Bioingegneria Via Ponzio 34/5

20133 Milan ITALY

andrea.aliverti@polimi.it AIMS

To better define and understand what do terms like mechanics, statics, kinematics and dynamics mean

To identify the fundamental variables describing respiratory mechanics To overview the different methods employed to measure these variables.

SUMMARY The mechanics of the respiratory system can be paralleled to that of any mechanical system. Pressure is equivalent to force, volume to position, flow to velocity and variation of flow to acceleration. The equation of motion relates these variable to resistive, elastic and inertial properties and it very useful to understand how applied forces produces motion and displacement of the system. Similarly to classical mechanics, which is subdivided into different aspects, respiratory mechanics can be described in terms of statics, dynamics and kinematics. Statics is the study of equilibrium (i.e., ‘position’ x of the system) and its relation to force, dynamics is the study of motion (i.e., velocity and acceleration) and its relation to force and kinematics deals with the implications of observed motions without regard for circumstances causing them. In the following sections, the main devices, their principles of measurement, and the most important problems associated with the different techniques are reviewed. Pressure measurement Pressure at the airways opening is measured at the side port of a mouthpiece by connecting a small tube to a pressure sensor. The measurement of pleural pressure definitely presents more difficulties, because accessing the pleural space is not feasible in human subjects. However, the pressure in the esophagus (Pes) is considered a useful surrogate for pleural pressure, because the esophageal lumen is subjected to essentially the same pressure swings as the pleural space. Measuring Pes requires the subject swallowing one balloon catheter, which consists of a thin-walled balloon a few cm in length sealed over the end of a thin plastic catheter typically about 100 cm in length. The catheter is passed through the nose after local anesthesia of the nasal mucosa and pharynx, and swallowed until the balloon reaches the lower third of the esophagus. A small volume of air (0.5-0.6 ml of air) is then injected through the catheter into the balloon so that there is free transmission of pressure between the interior of the balloon and the proximal end of the catheter, to which a pressure transducer is attached (providing a pressure measurement referred to atmospheric pressure). The volume of air in the balloon must be sufficient to prevent its walls from occluding the hole at the end of the catheter. On the other hand, the balloon volume must be low enough that its walls remain flaccid, or else there will be a transmural pressure gradient across the walls. Abdominal pressure (Pab) is usually measured with a similar technique based on a transducer-cathether-balloon systems. The catheter is still passed through the nose and swallowed until the balloon reaches the stomach, where gastric pressure (Pga) is

53

measured. Changes in Pga are generally considered good estimate of changes in Pab. For the measurement of Pga, measurements are made with a balloon volume of 1–2 ml. Regarding the transducers, in the respiratory laboratories, until relatively recently, the pressure measurements were based on variable reluctance sensors. These devices have now largely been replaced by the piezoresistive transducers, which have several advantages including an extremely high frequency response, robustness, and the fact that they can be manufactured using solid-state technology to be very small, light, and cheap. Flow measurement The basis of flow measurement in respiratory mechanics is the pneumotachograph, which is a calibrated resistance (R) across which a differential pressure is measured (Fig. 5). When gas flows through the pneumotachograph, there is a pressure drop (ΔP) from the upstream side of the resistance

to the downstream side that increases as flow (V’) increases. If R is independent of V’ over the range

of flows of interest, then the pneumotachograph is said to be linear. Linearity is generally achieved only within a range of flow values, and therefore it is necessary to calculate flow from a measurement of ΔP. The frequency response of a pneumotachograph depends on the construction of its resistive element, which may have a honeycomb arrangement of conduits or consist of a wire screen. The honeycomb type is less likely to become partially blocked by secretions but has a poorer frequency response than the screen type, as ΔP become dependent on a second term proportional to the derivative of the flow time through the inertia of the gas (which in turn is proportional to gas density and to the volume of the resistive element). Either type should be heated to above body temperature during prolonged use to avoid breath condensate from settling on the resistive element and changing its resistance (and hence altering the calibration of the device). Pneumotachographs can have a good frequency response above 20 Hz with a resonance occurring at around 70 Hz, provided that the associated differential transducer has a response at least that good and is connected with the shortest possible lengths of tubing. Although the resistive pneumotachograph is the principal device used to measure flow in respiratory applications, other devices have been proposed. Ultrasonic transducers based on differences in time-offlight of sound propagating into the direction of flow versus away from it have an excellent frequency response and avoid the problems of a resistive element becoming clogged with secretions. Devices based on the rate of cooling of a heated wire are also used. Volume measurement Nowadays, integration is performed digitally on a computer. The digitised flow signal consists of a series of data points separated by equal time intervals. A simple method for numerical integration is to calculate the area under the curve defined by the series of measured data. The key problem is that the sampling frequency should be high enough so that the errors involved in approximating the true curve between points are negligible. Another important problem is integration drift. When flow is integrated to yield volume, an upward or downward drift in the volume baseline is invariably seen. Some degree of drift is expected for purely physiological reasons such as carbon dioxide production/oxygen consumption ratio different than 1 and differences in water-vapour content between inspired and expired air). In addition, several methodological factors contribute to volume drift: a) temperature changes between inspired and expired gas; b) variations in temperature affecting the physical dimensions of the pneumotachograph due to the coefficients of thermal expansion of its components; c) changes in gas composition between inspiration and expiration and slight differences in the viscosities of the gas mixtures, with concomitant effects on the flows registered during inspiration and expiration by the pneumotachograph; d) leaks between the airway opening and the pneumotachograph, causing a discrepancy between the volume registered by the apparatus and that entering or leaving the lungs, and hence a drift in volume; e) zero offset in flow calibration; f) imperfections in the pneumotachometer response.

54

In principle, it might be possible to avoid drift in volume in this type of data record by pre-conditioning the inspired gas to body temperature pressure saturated (BTPS) conditions, continuously monitoring gas partial pressures in both the alveoli and the pulmonary arterial and venous blood to correct for respiratory-exchange ratios not equal to unity and eliminating all the methodological factors discussed above. However, this is extremely difficult, if not impossible, in practice. Consequently, it is never known how much of the baseline drift in volume is due to drift and how much represents a true change in absolute lung volume. The volume of gas entering the lungs can be measured directly with a spirometer attached to the mouth or from the pressure or flows emanating from a wholebody plethysmograph when the subject breathes through a conduit connected outside the plethysmograph. A more convenient but less accurate plethysmographic method is provided by the changes in trunk volume assessed with an inductance plethysmograph. Recently, a more accurate optical device, opto-electronic plethysmography (OEP), has been developed that allows detailed measurement of thoracic movement during breathing. The introduction of OEP solves the difficult problem of tracking absolute lung-volume changes on a breath-to-breath basis during almost any activity in which the chest wall can be visualised. OEP can also be combined with oesophageal and gastric pressure measurements (performed with standard balloon-catheter-transducer systems) to accurately quantify respiratory muscle dynamics. By compartmental volume analysis obtained with OEP and oesophageal and gastric pressure measurements, the pressures developed by the abdominal muscles and by the inspiratory and expiratory rib-cage muscles can be estimated as the difference between the dynamic compartmental pressure–volume loops and the corresponding relaxation curves of the abdomen and the rib cage, respectively. Conclusions

Studying respiratory mechanics means to study statics, kinematics and dynamics of the respiratory system and its components: airways, lung and chest wall.

The fundamental variables describing respiratory mechanics are flow, volume and pressure

Different methods are available to measure these variables, and each of them has to be

employed considering the inherent problems which are frequently forgotten. REFERENCES 1. Hedenstierna G. Respiratory Measurement. BMJ Books, 1998. 2. Aliverti A, Pedotti A (eds.). Mechanics of breathing: new insights from new technologies.

Springer Verlag, 2014. 3. Aliverti A. Monitoring of Respiratory Mechanics in the ICU: Models, Techniques and

Measurement Methods. In: Respiratory System and Artificial Ventilation .Lucangelo U, Pelosi P, Zin WA, Aliverti A (Eds.), Springer Verlag, 2007, pp. 73-97

55

EVALUATION 1. Dynamics refers to

a. The relationships between airflow and volume b. The relationships between pressure and flow c. The relationships between volume and pressure d. None of the above

2. The forced oscillation technique measures a. The action of the respiratory muscles b. The impedance of the respiratory system c. The pressure – flow relationship of the respiratory system at different lung volumes d. None of the above

3. When integrating airflow to obtain volume

a. Drift is present and is always positive b. Drift is present and is always negative c. Drift can be either positive or negative d. Drift is always negligible

4. Chest wall volume can be measured by a. Opto-electronic plethysmography b. Integration of airflow measured by a pneumotachometer c. Respiratory magnetometers d. None of the above

5. Volumetric flow rate is a. equal to flow velocity b. equal to density / time c. equal to flow velocity x area of the tube d. measured by an hot-wire anemometer

56

Measuring respiratory mechanics: invasive and noninvasive systems

Andrea Aliverti, PhD

Dipartimento di Elettronica, Informazione e Bioingegneria

Politecnico di Milano, Italy

57

Conflict of interest disclosure I have the following, real or perceived direct or indirect conflicts of interest that

relate to this presentation:

A.A. is one of the inventors of opto-electronic plethysmography. The patents are owned by the Politecnico di Milano (Polimi, Milan, his institution) and licensed to BTS Spa Company.

A.A. is one of the inventors of patents on a system for the automatic detection of the expiratory flow limitation by Forced Oscillation Technique (FOT). The patents are owned by Polimi and licensed to Philips Respironics and Restech, a spin-off Company of Polimi that develops and commercializes a FOT device. Both A.A. and Polimi own stocks of ResTech.

This event is accredited for CME credits by EBAP and speakers are required to disclose their potential conflict of interest going back 3 years prior to this presentation. The intent of this disclosure is not to prevent a speaker with a conflict of interest (any significant financial relationship a speaker has with manufacturers or providers of any commercial products or services relevant to the talk) from making a presentation, but rather to provide listeners with information on which they can make their own judgment. It remains for audience members to determine whether the speaker’s interests or relationships may influence the presentation.Drug or device advertisement is strictly forbidden.

58

Introduction

AIMS

• To better define and understand what do terms likemechanics, statics, kinematics and dynamics mean

• To identify the fundamental variables describingrespiratory mechanics

• To overview the different methods employed to measurethese variables

59

CLASSICAL MECHANICS IS DIVIDED INTO THREE MAIN BRANCHES

STATICS – the study of equilibrium (POSITION, x) and its relation to FORCE(s).

DYNAMICS – the study of MOTION (velocity and acceleration) and its relation to FORCE(s)

KINEMATICS – dealing with the implications of observed motions without regard for circumstances causing them

xF

),,( xxx

),( xxF

60

EQUATION OF MOTION

K

xMxBxKF

MF

xB

elasticforces

inertialforces

resistiveforces

61

RESPIRATORY VS ‘GENERIC’

MECHANICAL SYSTEM

Mechanical system Respiratory system

position, x volume, Vvelocity, x’ flow, V’

acceleration, x’’ var. flow, V’’

force, F pressure, P

rigidity, K elastance, E (distensibility, 1/K) (compliance, C)

friction, B resistance, R mass, M inertance, I

“state”

“properties”

62

EQUATION OF MOTIONFOR THE RESPIRATORY SYSTEM

VIVRVEP V

P

V’ E=1/C

R

63

RESPIRATORY MEASUREMENTS

Ventilation Diffusion Perfusion

Respiratory mechanics

Assessment of ‘active’

componentsAssessment of ‘passive’

components• respiratory muscle action

(force, work, power)• Mechanical properties

(Compliance, ResistanceInertance, Impedance)

Statics

Kinematics

Dynamics

64

RESPIRATORY MECHANICS: FUNDAMENTAL VARIABLES

Studying respiratory mechanics (i.e., statics, kinematicsand/or dynamics) requires the measurement of the following FUNDAMENTAL VARIABLES:

• Pressure (‘force’)

• Volume (‘position’)

• Flow (‘velocity’)

• Variation of Flow (‘acceleration’)

),,( xxxF

xxF

x

65

DIFFERENTIATION AND INTEGRATION

)(tx

dt

dxtx )(

dttxtx )()(

)(tx

differentiation

integration

66

OPERATING VOLUME = ‘POSITION’ OF THE SYSTEMFLOW = ‘MOTION’ OF THE SYSTEM

67

PRESSURE(S) = FORCE(S) ACTING ON THE SYSTEM

Pab

Ppl

Palv

Pao

PAW (Airways)

PL (Lung)

PDI (Diaphragm)

Pbs

PABW = Abdominal Wall

Location at which pressurescan be measuredor pressures of interest

68

THE SPIROGRAM IS KINEMATICS

69

FLOW-VOLUME LOOP KINEMATICS

70

MECHANISMS BEHIND FLOW-VOLUME CURVE DYNAMICS

(ISO-VOLUME PRESSURE FLOW-VOLUME CURVE)

Transpulmonary pressureflow

71

MECHANISMS BEHIND FLOW-VOLUME CURVE DYNAMICS

(ISO-VOLUME PRESSURE FLOW-VOLUME CURVE)

Gibson GJ, Clinical Tests of Respiratory Function, 2009

72

THORACO-ABDOMINAL VOLUME VARIATIONS KINEMATICS

rib cage

abdomen

chest wall

73

PRESSURE-VOLUME CURVES STATICS

Lung (L)Chest wall (CW)

Respiratory system(RS)

74

PRESSURE VOLUME CURVE STATICS

Loring et al, J Appl Physiol, 100:753-758, 2006

75

MAXIMAL PRESSURES (PIMAX AND PEMAX) STATICS

RespiratoryMuscle length

Insp.Mus.

length

Pressure

RespiratoryMuscle force

MaximalexpirationElastic recoil of

resp. System (Prs)

Pmus,e + Prs

Maximalinspiration

Pmus,i + Prs

PE max

PI max

FRC

TLC

RV

Exp.Mus.

length

flow

E=1/C

R

Lungvolume

50 cmH2O

76

INTERRUPTER TECHNIQUE DYNAMICS (first) AND STATICS (after)

77

ROHRER’S EQUATION (FLOW-DEPENDENTRESISTANCE) DYNAMICS

VVkVkPAW

21

JHT Bates. Lung Mechanics: An Inverse Modeling Approach, 2009

78

CLASSICAL STARLING EQUATIONPULMONARY INTERSTITIAL (FLUIDO) DYNAMICS

filtration flow from capillary to interstitium

Hydraulic pressure differencebetween plasma and interstitium

Colloid osmotic pressure difference between plasma and

interstitium

79

FORCED OSCILLATION TECHNIQUE MECHANICS

VIVRVEPPP rsrsrsbsaors

aors

rsrs

aors

rsrsao

VC

IjR

VCj

IjRP

1

1

80

CAMPBELL DIAGRAM MECHANICS

Loring et al, J Appl Physiol, 107: 309-314, 2009)

81

MEASUREMENTS OF FLOW / VELOCITY

Volumetric flow rate (also known as volume flow rate, rate of fluid flow or volume velocity)

volume of fluid which passes per unit time [m3/s in SI unit ]

𝑄 = 𝑉 =𝑑𝑉

𝑑𝑡=

𝑚

𝜌= 𝑣 ∙ 𝐴 (1)

where:v = flow velocity of the substance elementsA = cross-sectional vector area/surface 𝑚 = mass flow rate (kg/s) mass of a substance which passes per

unit of time [kg/s]. = density (kg/m3)

82

MEASUREMENTS OF FLOW / VELOCITY

• Equation (1) is only true for flat, plane cross-sections. In general, including curved surfaces, the equation becomes a surface integral:

𝑄 = 𝐴

𝑣 ∙ 𝑑𝐴

• The reason for the dot product is that the only volume flowing through the cross-section is the amount normal to the area; i.e., parallel to the unit normal. This amount is:

𝑄 = 𝑣 𝐴𝑐𝑜𝑠𝜃

83

MEASUREMENTS OF FLOW / VELOCITY

Mean velocity depends on fluid motion profile:

laminar motion, Re<2300

Vm = 0,5 Vmax

Turbolent motion, Re>10000

Vm ≈ 0,8-0,85 Vmax

84

FLOWMETERS TYPES

There are several different classifications of flowmeters.

In general, flowmeters can be divided into measurement systems for flow or velocity.

Another subdivision considers measurement systems with moving parts (dynamic) and with fixed parts (static)

We now briefly review the following types of flowmeters• Velocity flowmeters • Positive displacement meters• Variable area meters • Obstruction flowmeters

85

VELOCITY FLOWMETERS

Velocity flow measurement techniques allow for the measurement of total flow by measuring the velocity of the fluid within a fixed area duct or pipe. The technique uses a measuring probe to determine the velocity of the fluid, usually in the center portion of the pipe. To provide reasonably accurate results, the velocity measurement of the flow must be made well within the duct, to minimize the effects of the boundary layers. For this reason ducts of small diameter typically do not fair well with this technique. The technique also requires to be in a laminar flow environment. The results in a turbulent flow area suffer in stability and accuracy.

Hot-wire anemometer

Pitot tube

86

MOVING MEMBER/VARIABLE AREAFLOWMETERS

Axial Turbine Flowmeter Radial Turbine Flowmeter

Rotameter

87

OBSTRUCTION FLOW METERS

• Venturi Flow meter

• Orifice plate

• Nozzle

88

PNEUMOTACHOMETER (PNT)

The flow sensors that historically have been, and continue to be, the mainstay of the respiratory laboratory utilize flow resistors with approximately linear pressure–flow relationships. These devices are usually referred to as ‘pneumotachometers’

a term that, in general, is synonymous with gas volume flowmeter. Flow-resistance pneumotachometers are easy to use and can distinguish the directions of alternating flows. They also have sufficient accuracy, sensitivity, linearity, and frequency response for most clinical applications. In addition, they use the same differential pressure sensors and amplifiers required for other respiratory measurements.

89

PNEUMOTACHOMETER (PNT)

QLQRP

2r

lL

44

1288

d

l

r

lR

QRP

l, r, d = length, radius and diameter of the channels = dynamic viscosity of the fluid = density of the fluid

90

DIFFERENTIATION AND INTEGRATION

)(tx

dt

dxtx )(

dttxtx )()(

)(tx

differentiation

integration

91

FLOW VOLUME

• Volume and flow are related each other by time integration (flow volume) or differentiation (volume flow).

• Typically, the primary measured respiratory signal is flow. This must be integrated with respect to time to produce volume; this is most conveniently performed using a computer.

92

FLOW VOLUME

There are a variety of numerical integration methods available. One of the simplest is the so-called trapezoidal rule

Bates et al, Eur Respir J 2000; 16: 1180±1192

93

THE PROBLEM OF VOLUME DRIFT

offsetshift

drift

Volume dV / dt · dtFlow dV / dt

94

CORRECTION OF VOLUME DRIFT IS CRITICAL

95

SOURCES OF INTEGRATION DRIFT

• Temperature changes between inspired and expired gasIf inspired air is not warmed to body temperature before passing through the PNT, it has a different viscosity and density to expired air, • Changes in gas composition between inspiration andexpiration• Leaks• Imperfections in the pneumotachometer response• Zero offset in flow calibration

96

Avoiding vapor condensation in PNTs is particularly important. In fact, the small tubes (in Fleisch-type) and holes (in mesh-type PNT) can be easily closed by water.

As a consequence, the cross sectional area of the resistive element diminishes and flow resistance increases.

Furthermore, when water vapor condensates, a change in the gax mixture occurs.

97

To prevent these problems, particularly when a series of consecutive breaths must be studied the resistance of the PNT is heated.

The PNT is commonly equipped with an heater constituted by an electrical resistance.

The heating system can be within (wires) or around (radiant ribbons) the conduit where airflow passes.

98

LUNG VOLUME MEASUREMENTS

• relative measurements (variations of alveolar gas)- spirometry

- flow integration

• absolute measurements (alveolar gas volume)- washout techniques

(e.g. N2)

- dilution techniques

(e.g. He)

- total body plethysmography

West, Pulmonary Physiology

99

‘ABSOLUTE’ VOLUME MEASUREMENTS

Dilution methods (e.g. He)Total body plethysmography

Washout methods (e.g. N2)

nitrogenanalyzer

one-way valve

Stop-cock valve

PNT+integrator

100

LUNG AND SPIROMETER ARE IN DIFFERENT CONDITIONS

Spirometer

N Q VAWO AWO AWO L L

Lung

Molar flow of gas at airway opening AWON

)(

)(

flowvolumetricgasVQ

densitymolarV

N

molesgasofnumberN

volumeV

L

t

t

LLL

t

AWOL

AWO V)t(V)t(VdtVdtQt

00

0

101

2 L

~1.8 L

T= 37°P=Patm - PH2O,37°= Patm – 47 mmHg

ATPS (Ambient Temperature and Pressure,

saturated)

BTPS(Body Temperature and Pressure,

saturated)

T= 20°P=Patm - PH2O,20°= Patm – 17.5 mmHg

102

CHEST WALL DISPLACEMENT/VOLUME MEASUREMENTS

“Chest wall = all parts of the body outside the lung which share changes in the volume of the lungs“ (Konno and Mead, J Appl Physiol, 22:407-422, 1967)

During breathing, chest wall varies not onlyvolume, but also shape Measurement has to be done in severalpoints of the thoraco-abdominal wall

• Where ?• How any “degrees of freedom” does chestwall have during breathing?

103

CHEST WALL DISPLACEMENT/VOLUMEMEASUREMENTS

RC

AB

magnetometers Inductiveplethysmography

Opto-electronic plethysmography

Structured light plethysmography

Wearable resistive/fiber opticsensors

104

CHEST WALL VOLUMES

Vrc,p(liters)

10.4

10.5

10.6

10.7

10.8

10.9

11.0

11.1

Vrc,a(liters)

2.6

2.8

Vab(liters)

7.0

7.2

7.4

7.6

Gauss theorem

Abdomen (AB)

AbdominalRib Cage

(RCa)

r r rF dV F ndS

SV

PulmonaryRib cage

(RCp)

105

OPERATING VOLUMES DURINGEXERCISE

106

CHEST WALL MECHANICS

107

CHEST WALL MECHANICS

Pab

Ppl

Palv

Pao

AW

L

DI

Pbs

AB

Pab

Ppl

Palv

Pao

Pbs

AW

L

DI

ABRC

chestwall

diaphragm

Abdom.muscles

ribcagemuscles

108

MEASUREMENT OF ESOPHAGEALAND GASTRIC PRESSURES

109

54

RESPIRATORY MUSCLE MECHANICS

PaoPesPga

Vrc,pVrc,aVab

PRESSURES + VOLUMES

MECHANICS

110

SPONTANEOUS BREATHING AT REST

Chest wall Volume(Vcw, liters)

Esophageal Pressure (Pes, cmH2O)

Airflow(L/sec)

Gastric Pressure (Pga, cmH2O)

Time (sec)

111

Slow inspiration(“quasi-static”)

Chest wall“relaxation curve”

Vcw

Pes

Flow

Time (sec)

112

RESPIRATORY MECHANICS

113

TRANSDIAPHRAGMATIC PRESSUREPDI = PES-PGA

Chest wall Volume(Vcw, liters)

TransdiaphragmaticPressure

(Pdi, cmH2O)

Airflow(L/sec)

Time (sec)

114

59

PRESSURE DEVELOPED BY RIB CAGE (PRCM) AND ABDOMINAL (PABM) MUSCLES

Aliverti, J Appl Physiol , 83:1256-1269, 1997

115

Conclusion

• Studying respiratory mechanics means to study statics, kinematics and dynamics of the respiratory system and its components: airways, lung and chest wall

• The fundamental variables describing respiratorymechanics are flow, volume and pressure

• different methods are available to measure thesevariables, and each of them has to be employedconsidering the inherent problems which are frequentlyforgotten

116

Using respiratory mechanics to explain and interpret spirometry

Prof. John Gibson Emeritus Professor of Respiratory Medicine

Newcastle University Newcastle upon Tyne UNITED KINGDOM

john.gibson@ncl.ac.uk

SUMMARY The apparent simplicity of the forced expiratory spirogram (volume v. time) is deceptive as the underlying respiratory mechanics are complex. The factors which determine maximum expiratory flow from the lungs include (1) those related to the airway itself: the complex geometry and overall calibre of different airway generations, together with the dynamic mechanical properties of the airway walls, both intrinsic and extrinsic; (2) the pressure driving expiratory airflow: a function of the strength and velocity of shortening of the expiratory muscles, together with the lung recoil pressure and (3) the density and viscosity of the air being expired. The pleural pressure (Ppl) resulting from expiratory muscle contraction not only increases the driving pressure for expiration, but also potentially compresses the central extrapulmonary airways upstream of the thoracic inlet. Because the airways are compliant to a varying degree, the effect of increasing expiratory effort beyond a certain level may be negated by progressive narrowing. A major mechanism underlying dynamic compression of the airways and consequent expiratory flow limitation is the Bernoulli phenomenon. In accordance with this, the velocity of air (or any fluid) flowing from a wider to a narrower tube has to increase to accommodate a constant flow. This applies to the airways during expiration as the total cross-sectional area decreases dramatically from the alveoli to the trachea. The consequent “convective acceleration” of the expired air is accompanied by a pressure drop within the airway, increasing the pressure difference across the airway wall, which, in turn, can produce a “suction” effect

narrowing the airway further to a level determined by its compliance. An alternative explanation for expiratory flow limitation relates to the speed with which a pressure wave is conducted along the airway wall; in a completely rigid tube this would be at the speed of sound, but, in a compliant airway, air added to the system by alveolar deflation is accommodated by bulging of the airway wall as a wave, the velocity of which depends on the mechanical properties of the airway. According to this concept, flow becomes limited when the local air velocity reaches the velocity of the pressure wave. The consequences of dynamic narrowing of the airways during forced expiration are most easily understood using the classical equal pressure point (EPP) analysis of Mead et al. At most lung volumes, as expiratory effort is increased, points are reached in the airways at which intra- and extra- bronchial pressures become equal; downstream (ie on the mouth side) of these EPPs there is a tendency for dynamic compression and flow limitation to occur. Evidence indicates that in normal individuals EPPs occur in lobar or segmental bronchi over much of the vital capacity but they move upstream (nearer to the alveoli) at low volumes and the smaller airways then become more important in determining maximum expiratory flow. (In this simple model the EPPs and sites of flow limitation are not necessarily identical as the contribution of airway elasticity itself is not considered). In the EPP model, the only factors determining flow during forced expiration are the resistance of the “upstream” airways and the pressure driving flow along this segment ie

the difference between alveolar and pleural pressures, which is the lung recoil pressure at the relevant volume. If an airway is compressed sufficiently that it closes completely the pressure immediately upstream equalises with the driving alveolar pressure, creating a positive (outward) transmural pressure which then

117

tends to reopen the airway. In effect, therefore, due to this instability airways may oscillate, with the resulting vibration contributing to the wheezing heard on forced expiration, even in healthy subjects. The effects of forced expiration and dynamic compression can be visualised using the isovolume pressure flow (IVPF) curves described by Fry and Hyatt. A curve can be constructed for any lung volume by measuring alveolar pressure (eg by oesophageal manometry to estimate pleural pressure) during a series of varying efforts and relating flow to alveolar pressure (sum of Ppl and lung recoil pressure). At most lung volumes below 10-20% FVC expired, the expiratory IVPF curve shows an initial rise followed by progressively smaller increases in flow until it reaches a plateau which represents the maximum value attainable at that lung volume (equivalent to the value recorded at that volume during a continuous forced expiration). At high lung volumes the IVPF relationship is curvilinear, but without a plateau ie. maximum flow at these volumes (including the volume at which peak expiratory flow is measured) is not flow-limited, but effort- dependent because the maximum pressure and the velocity of shortening of the expiratory muscles under dynamic conditions are insufficient for flow limitation to be established. On inspiration IVPF curves do not show plateaux as intra-airway pressure is greater than extra-airway pressure at all lung volumes and the normal tendency is for some widening rather than compression of the airway. Consequently, maximum inspiratory flow is completely effort-dependent, varying only with the lung volume and the dynamic performance of the inspiratory muscles. The flow rates recorded on the much more familiar maximum expiratory (MEFV) and inspiratory (MIFV) flow-volume curves, obtained during single maximum efforts from TLC and RV, represent the maxima on IVPF curves. In diffuse intrathoracic airway obstruction (eg COPD or asthma), IVPF curves show the expected reductions in flow for a given effort and, compared to the normal subject, they show relatively more independence of effort as the maximum alveolar pressure necessary to produce maximum expiratory flow is generally less than in healthy individuals. MEFV curves in COPD or asthma show a characteristic concavity to the volume axis due to asynchrony of lung “emptying”. In qualitative terms the appearance of the MEFV curve is an exaggeration of the pattern seen with healthy ageing. The so-called “airway collapse” MEFV pattern, with a very sharp peak followed

by a dramatic fall in maximum expiratory flow is determined by the severity of airway narrowing; it is not specific for emphysema as it is also seen with severe airway obstruction due to asthma. An important factor which is often ignored during forced expiration is the effect of thoracic gas compression due to the high alveolar pressures which can be generated; such pressures reduce the volume of thoracic gas (Boyle’s law) and therefore reduce lung recoil pressure. This effect is greater in patients with more severe airway obstruction, especially in the presence of a markedly raised RV. It can be demonstrated on the expiratory IVPF curve if “isovolume” is defined in terns of %FVC expired rather than in terms of thoracic gas volume (TGV). Use of expired volume, therefore, may give a spurious impression of “negative effort-dependence” (ie an apparent decline in expiratory flow with efforts exceeding that necessary to generate

maximum flow). The same phenomenon is responsible for the frequent observation that many patients deliver a greater “FEV1” with a less than maximal expiratory effort. The “compression artefact” is also one of the factors responsible for the observation that tidal expiratory flow may apparently exceed maximal at the same lung volume if MEFV and tidal FV curves are superimposed. However, this is not the only factor as non-homogeneity of lung emptying and differences in volume- and time- history between tidal and forced expiration also contribute. The MEFV and volume-time (spirogram) curves each contain the same information but each displays only two of the related variables: flow, volume and time. The MEFV curve allows better visualisation of the early part of expiration and is therefore useful for assessing adequacy of effort as flows in early expiration are the most effort-dependent; the spirogram optimises visualisation of late expiration, when the effect of very slowly emptying alveoli (with long time constants) is dominant.

118

In clinical practice, maximum flow-volume curves are particularly useful for recognition and categorisation of localised narrowing of the central airway, whether extra- or intra-thoracic, with typical patterns well recognised. Extrathoracic narrowing is easily modelled by having a healthy subject perform maximum inspiration and expiration through an added resistance at the mouth. The maximum flows most affected are those which are normally most effort-dependent ie in early expiration and throughout forceful inspiration (the precise converse of the pattern seen with the diffuse intrathoracic narrowing of COPD or asthma). With relatively fixed narrowing, maximum inspiratory and expiratory flows are similarly affected, while a markedly disproportionate reduction in maximum inspiratory flow indicates appreciable dynamic narrowing. This is seen with a highly compliant upper airway (eg due to tracheal damage) or a “ball valve” type

phenomenon, as occurs occasionally with a mobile tumour which is drawn into the inspiratory airstream. Upper airway narrowing is less easily visualised on the volume-time display, but the characteristic “straight

spirogram” associated with a plateau of flow in early expiration may be detectable. Unilateral main stem bronchus obstruction should be considered if the MIFV curve shows markedly delayed filling of the lungs with a “parallelogram” appearance and a “biphasic” spirogram typical of two-compartment emptying. REFERENCES General 1. Fry DL, Hyatt RE. Pulmonary mechanics: a unified analysis of the relationship between pressure,

volume and gas flow in the lungs of normal and diseased human subjects. Am J Med 1960; 29: 672-89

2. Mead J, Turner JM, Macklem PT, Little JB. Significance of the relationship between lung recoil and maximum flow. J Appl Physiol 1967; 22: 95-108

3. Dawson SV, Elliott EA. Wave-speed limitation on expiratory flow – a unifying concept. J Appl Physiol 1977; 43: 498-515

4. Mead J. Expiratory flow limitation: a physiologist’s point of view. Federation Proc 1980; 39: 2771-5

5. Hyatt RE. Forced expiration. In Macklem PT, Mead J (eds) Handbook of Physiology section 3 vol III, American Physiological Society 1986; 295-314

6. Pride NB, Macklem PT. Lung mechanics in disease. In Macklem PT, Mead J (eds) Handbook of Physiology section 3 vol III, American Physiological Society 1986; 659-92.

7. Pedersen OF, Butler JP. Expiratory flow limitation. Comp Physiol 2011; 1: 1861-82 8. Gibson GJ. Clinical tests of respiratory function. 3rd ed Hodder Arnold 2009

Thoracic gas compression 9. Mead J. Volume-displacement body plethysmograph for respiratory measurements in human

subjects. J Appl Physiol 1960; 15: 736-40 10. Ingram RH, Schilder DP. Effect of gas compression on pulmonary pressure, flow and volume

relationship. J Appl Physiol 1966; 21: 1821-6 11. Krowka MJ, Enright PL, Rodarte JR, Hyatt RE. Effect of effort on measurement of forced

expiratory volume in one second. Am Rev Respir Dis 1987; 136: 829-33 COPD / asthma 12. Macklem PT, Fraser RG, Brown WG. Bronchial pressure measurements in emphysema and

bronchitis. J Clin Invest 1965; 44: 897-905 13. Leaver DG, Tattersfield AE, Pride NB. Contributions of loss of lung recoil and of enhanced

airways collapsibility to the airflow obstruction of chronic btronchitis and emphysema. J Clin Invest 1973: 52: 2117-28

14. Jayamanne DS, Epstein H, Goldring RM. Flow-volume contour in COPD: correlation with pulmonary mechanics. Chest 1980; 77: 749-57

119

15. Hogg JC. Pathophysiology of airflow limitation in chronic obstructive pulmonary disease. Lancet 2004; 364: 709-21

Upper/Central airway obstruction 16. Miller RD, Hyatt RE. Evaluation of obstructing lesions of the trachea and larynx by flow-volume

curves. Am Rev Respir Dis 1973; 108: 476-81 17. Gibson GJ, Pride NB, Empey DW. The role of inspiratory dynamic compression in upper airway

obstruction. Am Rev Respir Dis 1973; 108: 1352-60 18. Gascoigne AD, Corris PA, Dark JD, Gibson GJ. The biphasic spirogram: a clue to unilateral

narrowing of a mainstem bronchus. Thorax 1990; 45: 637-8 EVALUATION

1. During forced expiration:

a. lung recoil pressure exceeds alveolar pressure b. intraairway pressure at points of flow limitation exceeds alveolar pressure c. according to the Bernoulli theorem, flow velocity decreases as air passes from a wider

to a narrower airway d. equal pressure points (EPPs) in the airways may differ from the sites of maximum

dynamic narrowing e. audible wheezing implies abnormal airway narrowing.

2. The isovolume pressure-flow (IVPF) curve:

a. can be obtained during a single forced expiratory manoeuvre b. has a slope equivalent to airway resistance c. in normal subjects shows effort-dependence at low lung volumes eg 75% FVC expired d. shows the minimum alveolar pressure needed to generate maximum flow at a given

lung volume e. in COPD shows a plateau of flow on inspiration

3. In patients with COPD:

a. maximum flow volume curves show greater reduction of inspiratory than expiratory flow

b. thoracic gas compression during forced expiration has a greater effect on measurements of FEV1 than in normal subjects.

c. Maximum flow-volume curves help to distinguish the condition from asthma d. concavity of the descending limb of the maximum flow-volume curve (ie convexity

towards the volume axis) is diagnostic of airway obstruction. e. lung recoil pressure is usually increased

4. In patients with extrathoracic airway obstruction:

a. maximum expiratory flow is usually reduced more than maximum inspiratory flow b. a plateau of maximum expiratory flow towards the end of forced expiration is

characteristic c. intrathoracic flow limitation during forced expiration does not occur d. peak expiratory flow is typically reduced by a greater proportion than FEV1 e. the intensity of stridor indicates the severity of airway narrowing

120

Using respiratory mechanics to explain and interpret spirometry

John Gibson Emeritus Professor of Respiratory Medicine

Newcastle University UK

121

Conflict of interest disclosure

I have no real or perceived, direct or indirect

conflicts of interest that relate to this presentation.

122

AIMS

1. To understand the interrelationships of pressure, airflow, volume and time during forced expiration and the various factors which influence these variables.

2. To understand the classical Equal Pressure Point model of forced expiration

3 To interpret abnormalities of maximum inspiratory and expiratory flow in airway disease

123

Factors Determining Maximum Expiratory Flow

Airway: geometry calibre* compliance (intrinsic, extrinsic)

Driving pressure: muscle force*, velocity lung recoil pressure*

Gas: density viscosity

* varies with lung volume

124

Intrathoracic Pressures

alveolar pressure = pleural pressure + lung recoil pressure

Palv = Ppl + Pel

el

125

Intrathoracic airway surrounded by pleural pressure

126

Bernoulli Principle

convective acceleration: velocity ↑, pressure ↓

as tube narrows

Daniel Bernoulli

127

Expiratory flow limitation: Bernoulli mechanism

↓ X-sectional area

suction effect ↑linear velocity

(convective acceleration)

↑transmural pressure ↓intra-airway pressure

128

Expiratory flow limitation: wave speed mechanism

In a rigid tube a pressure pulse travels at the speed of sound

In a compliant airway the air added is accommodated by bulging of the wall as a wave; wave velocity is dependent on the mechanical properties of the airway wall

Flow is limited when local air velocity reaches the wave velocity (“choke pont”)

129

Equal Pressure Point (EPP) model

80

upstream downstream segment

after Mead et al 1967

130

What happens if the airway closes?

80 80

pressure immediately upstream exceeds pleural pressure, tending to reopen the airway

131

Isovolume Pressure Flow (IVPF) Curve(Fry and Hyatt 1960)

Relates the driving pressure to expiratory flow at a specific lung volume

Constructed by measuring alveolar pressure and flow at a specific lung volume during several expiratory

efforts of varying intensity

Not used for clinical purposes but helps understanding of forced expiration

132

Expiratory Isovolume Pressure–Flow (IVPF)curve at 50% FVC

133

Expiratory and Inspiratory Isovolume Pressure–Flow (IVPF) curves at 50% FVC

134

Expiratory IVPF curve at 50%FVC(1) relaxed expiration

1

1

135

Expiratory IVPF curve at 50%FVC(2) modest effort

1

2

1

2

136

Expiratory IVPF curve at 50%FVC (3) maximum effort → no further increase in flow

1

2 3

1

2

3

137

Relation of IVPF and MEFV curves (after Fry and Hyatt 1960):

(a) 10% FVC expired/peak flow

IVPF MEFV

138

Relation of IVPF and MEFV curves (after Fry and Hyatt 1960):

(b) 25% FVC expired

IVPF MEFV

139

Relation of IVPF and MEFV curves (after Fry and Hyatt 1960):

(c) 50% FVC expired

IVPF MEFV

140

Relation of IVPF and MEFV curves (after Fry and Hyatt 1960):

(d) 75% FVC expired

IVPF MEFV

141

Relation of IVPF and MEFV curves (after Fry and Hyatt 1960):

complete FVC

142

IVPF curves at various lung volumes: expiratory and inspiratory

a 10% FVC exp

b 25% FVC exp

c 50% FVC exp.

d 75% FVC exp.

143

Relation of IVPF and maximum flow-volume curves:expiratory and inspiratory

IVPF MEFV/MIFV

144

IVPF curves at 50% FVC:normal v. COPD

normal

COPD

145

Typical MEFV/MIFV curves

normal COPD

146

volume

FV curves in COPD are an exaggeration of normal ageing

normal young normal older COPD/asthma

147

Effect of unequal time constants on shape of MEFV curve

148

Measuring volume change during forced expiration

expired volume variable volume plethysmograph

after Mead 1960

149

Thoracic gas compression during forced expiration

Inevitable (Boyle’s Law: P x V = constant)

Effect greater with larger lungs (especially large RV)

eg if TLC = 10 litres

FVC = 2 litres

Palv = 100 cm H2O (0.1 atmos)

thoracic gas is compressed by 10%

therefore lung volume is reduced to 9 litres or 50% FVC

150

Effect of thoracic gas compression on“isovolume” pressure-flow curve

TGV 50%FVC

expired volume 50% FVC

151

Thoracic gas compression during forced expiration(after Ingram and Schilder 1976)

volume expired

TGV

declining TGV is due to (1) volume expired + (2) volume compressed

152

10 patients with mild-moderate COPD (mean FEV1 77% predicted)

and variable FEV1 results:

FEV1 from efforts with the highest PEF < largest FEV1 in all

mean difference 200 ml, range 160 – 350 ml.

Am Rev Respir Dis 1987; 136: 829

Implications of thoracic gas compression for FEV1 measurements

153

Summary

1. Dynamic compression during forced expiration results from Bernoulli effects and limitation by wave speed.

2. Maximum expiratory flow is relatively effort-independent over the lower 75 to 85% of the FVC, provided that the subject has inspired fully and exerts a modest effort

3. Thoracic gas compression may reduce maximum expiratory flow and FEV1 particularly in patients with severe hyperinflation.

4. The abnormalities of Flow Volume curves in diffuse airway obstruction are essentially an exaggeration of the changes seen with normal ageing.

5. Extrathoracic airway obstruction impairs those maximal flows which are most dependent on effort.

154

Complementary methods of visualising forced expiration

flow v. volume volume v. time

time missing (marker for FEV1) flow missing (tangent to the curve)

optimal for early expiration (effort) shows late expiration (slow emptying)

155

Extrathoracic (Upper) airway obstruction

the extra resistance is in series with the normal airway resistance

effect can be mimicked by adding an external resistance at the mouth156

IVPF and MEFV: 25% (b) and 75% (d) FVC expired – normal

157

Effect of UAWO at 25% FVC expired

158

Effect of UAWO at 25% (b) and 75% (d) expired

159

Effect of UAWO at 25% (b) and 75% (d) expired

160

Dynamic inspiratory narrowing of trachea

within lesion below lesion pedunculated tumour

161

Maximum Flow-Volume curves: pronounced dynamic inspiratory narrowing

162

Typical Flow-Volume curves in central/upper airway obstruction

volume volume

VEmax

VImax

VEmax

VImax

extrathoracicintrathoracic

unilateralmainstem

“saw tooth”

163

Using respiratory mechanics to study exercise limitation in disease

Dr Lee Romer College of Health and Life Sciences

Brunel University UB8 3PH Uxbridge

UNITED KINGDOM lee.romer@brunel.ac.uk

164

Using respiratory mechanics to study exercise limitation in disease

Lee Romer, PhD

College of Health and Life SciencesBrunel University London

UK

165

Conflict of interest disclosure I have no, real or perceived, direct or indirect conflicts of interest that relate to

this presentation. I have the following, real or perceived direct or indirect conflicts of interest that

relate to this presentation: Affiliation / financial interest Nature of conflict / commercial company name

Tobacco-industry and tobacco corporate affiliate relatedconflict of interest

Grants/research support (to myself, my institution or department):

Honoraria or consultation fees:

Participation in a company sponsored bureau:

Stock shareholder:

Spouse/partner:

Other support or other potential conflict of interest:

This event is accredited for CME credits by EBAP and speakers are required to disclose their potential conflict of interest going back 3 years prior to this presentation. The intent of this disclosure is not to prevent a speaker with a conflict of interest (any significant financial relationship a speaker has with manufacturers or providers of any commercial products or services relevant to the talk) from making a presentation, but rather to provide listeners with information on which they can make their own judgment. It remains for audience members to determine whether the speaker’s interests or relationships may influence the presentation.Drug or device advertisement is strictly forbidden.

X

166

AIMS

• Review the pressure-flow-volume relationships during exercise in health and disease

• Identify the respiratory influences on O2 transport, with specific focus on respiratory muscle work and fatigue

• Discuss how respiratory influences might exacerbate exercise-induced fatigue and compromise endurance performance

167

Exercise-induced respiratory muscle work/fatigue

Exercise-induced arterial hypoxaemia

Muscle O2 transport =

limb blood flow (QL) arterial O2 content (CaO2).

OVERVIEW

168

Exercise-induced respiratory muscle work/fatigue

Exercise-induced arterial hypoxaemia

Muscle O2 transport =

limb blood flow (QL) arterial O2 content (CaO2).

OVERVIEW

169

Exercise-induced arterial hypoxaemia

Muscle O2 transport =

limb blood flow (QL) arterial O2 content (CaO2)

OVERVIEW

.

Exercise-induced respiratory muscle work/fatigue

170

RESPIRATORY MUSCLES

171

EELV = TLC – IC

EILV = EELV + VT

FLOW-VOLUME RELATIONSHIPS AT REST AND DURING EXERCISE

NORMALS

172

IC IC IC

EILV

ERV

Rest Sub-maximal exercise Peak exercise

RV

TLC

IRV

PFT

EELV

LUNG VOLUME RESPONSE TO EXERCISE IN NORMALS

173

NORMALS

FLOW-VOLUME RELATIONSHIPS AT REST AND DURING EXERCISE

COPD

174

RV

TLC

LUNG VOLUME RESPONSE TO EXERCISE IN COPD

Rest Sub-maximal exercise Peak exercisePFT

175

RV

TLC

RV

TLC

Dynamic HyperinflationFRC

Static Hyperinflation

LUNG VOLUME RESPONSE TO EXERCISE IN COPD

Rest Sub-maximal exercise Peak exercisePFT

176

PRESSURE-VOLUME RELATIONSHIPS AT REST AND DURING EXERCISE

NORMALS COPD

177

Johnson et al., 1991, 1992, J Appl Physiol

Fit 25 yr old

Ventilation = 169 L/min

Fit 70 yr old

Ventilation = 109 L/min

COPD 70 yr old

Ventilation = 64 L/min

PRESSURE-VOLUME RELATIONSHIPS AT REST AND DURING EXERCISE

178

O2 COST OF EXERCISE HYPERPNOEA

179

Levison & Cherniack, 1968, J Appl Physiol

COPD

NORMAL

O2 COST OF EXERCISE HYPERPNOEA IN COPD

180

DO THE RESPIRATORY MUSCLES FATIGUE IN RESPONSE TO EXERCISE?

181

ASSESSING DIAPHRAGM FATIGUE

182

ASSESSING DIAPHRAGM FATIGUE

183

Johnson et al., 1993, J Physiol

EXERCISE-INDUCED DIAPHRAGM FATIGUE IN NORMALS

Pdi 20-30% (p < 0.01)

184

Bachasson et al., 2013, PLOS ONE

EXERCISE-INDUCED DIAPHRAGM FATIGUE IN COPD

Pdi,tw 20% (p < 0.001)

6/11 patients >15% Pdi,tw

185

ASSESSING ABDOMINAL MUSCLE FATIGUE

186

Taylor et al., 2006, J Appl Physiol

Pga 25% (p <0.01)

EXERCISE-INDUCED ABDOMINAL MUSCLE FATIGUE IN NORMALS

187

EXERCISE-INDUCED ABDOMINAL MUSCLE FATIGUE IN COPD

3/11 patients >15% Pga,tw

Pga,tw ~6% (p > 0.05)

Bachasson et al., 2013, PLOS ONE188

EXERCISE-INDUCED QUADRICEPS MUSCLE FATIGUE IN COPD

Qtw ~34% (p < 0.001)

13/15 patients >15% Qtw

Bachasson et al., 2013, PLOS ONE189

HOW DOES FATIGUING RESPIRATORY MUSCLE WORK IMPAIR EXERCISE PERFORMANCE?

190

RESPIRATORY MUSCLE METABOREFLEX

- Fatiguing contractions of the diaphragm, expiratory and accessory respiratory muscles

- Reflex activating metabolites- Group III/IV phrenic afferent discharge

- Sympathetic efferent discharge- Limb vasoconstriction- O2 transport

191

-50

-40

-30

-20

-10

0

10

0 200 400 600 800 1000

Diaphragm Work

Dia

phra

gm F

atig

ue

(% C

hang

e in

Pdi

)

Rest Hyperpnoea

Exercise Hyperpnoea

SEVERITY OF RESPIRATORY MUSCLE FATIGUE: REST VS. EXERCISE

Babcock et al., 1995, J Appl Physiol192

COMPETITION FOR BLOOD FLOW BETWEEN RESPIRATORY AND LOCOMOTOR MUSCLES

Harms et al., 1997, J Appl Physiol193

Harms et al., 1997, J Appl Physiol

COMPETITION FOR BLOOD FLOW BETWEEN RESPIRATORY AND LOCOMOTOR MUSCLES

194

Inspiratory unload

(13.2 ± 0.9 min)

INFLUENCE OF INSPIRATORY MUSCLE WORK ON LIMB MUSCLE FATIGUE

Romer et al., 2006, J Physiol195

Plateau No plateau

COMPETITION FOR BLOOD FLOW BETWEEN RESPIRATORY AND LOCOMOTOR MUSCLES IN COPD?

Simon et al., 2001, J Appl Physiol196

-10

-20

% twitch force(Pre vs Post exercise)

(62 W, 10 ± 1 min)Control

87%SaO2 (%)

limbs

dyspnoea

∫Poe (% of control)

8.5

100%

100%

8.7RPE

PAV+HeO2

92%

40%

70%

5.4

4.9

*

[p < .05]

.60 FIO2

100%

90%

75%

5.4

4.9

*

Insp

Exp

• Similar 28 to 37% in twitch force via WOB &/or SaO2

• Most of peripheral fatigue still remains

• inherent muscle fatigability is most important

Amann et al., 2010, Am J Physiol

EFFECTS OF RESPIRATORY MUSCLE UNLOADING &/OR PREVENTING O2 DESATURATION IN COPD

197

SUMMARY

198

CONCLUSION

• Three potential routes inhibit central command of locomotor muscles and limit exercise in COPD

• Underlying common events are hyperinflation and increased work of breathing produced by high ventilatory requirements and expiratory flow limitation

• Dyspnoeic sensations are a major cause of inhibition of central command of locomotion

• High expiratory intrathoracic pressures or respiratory muscle metabolite accumulation precipitate, respectively, a reduced stroke volume and/or a reflex-mediated sympathetic vasoconstriction of locomotor muscle vasculature

• Both routes would increase feedback via III-IV afferents to increase the drive to breathe and inhibit central command to locomotor muscles, precipitating exercise limitation

199

Additional course resources Readings and guidelines 1. Bates, J.H., et al., Oscillation mechanics of the respiratory system. Compr Physiol, 2011. 1(3): p.

1233-72. 2. Bates, J.H.T., Lung mechanics: An inverse modeling approach. 2009, Cambridge: Cambridge

University Press. 3. Lauzon, A.M., et al., A multi-scale approach to airway hyperresponsiveness: from molecule to

organ. Front Physiol, 2012. 3: p. 191.

200

Faculty disclosures

Prof. Andrea Aliverti is one of the inventors of opto-electronic plethysmography. The patents are owned by the Politecnico di Milano (Polimi, Milan) and licensed to BTS Spa Company. He is also one of the inventors of patents on a system for the automatic detection of the expiratory flow limitation by Forced Oscillation Technique (FOT). The patents are owned by Polimi and licensed to Philips Respironics and Restech, a spin-off Company of Polimi that develops and commercializes a FOT device. Both Prof. Aliverti and Polimi own stocks of ResTech.

201

Faculty contact information Prof. Andrea Aliverti Politecnico di Milano Dipartimento di Elettronica, Informazione e Bioingegneria Via Ponzio 34/5 20133 Milan ITALY andrea.aliverti@polimi.it Dr Jason H.T. Bates University of Vermont College of Medicine Burlington Vermont 05405-007 UNITED STATES OF AMERICA jason.h.bates@uvm.edu Prof. John Gibson Emeritus Professor of Respiratory Medicine Newcastle University Newcastle upon Tyne UNITED KINGDOM john.gibson@ncl.ac.uk

Prof. Michael Polkey Royal Brompton Hospital Sydney Street SW3 6NP London UNITED KINGDOM m.polkey@rbh.nthames.nhs.uk Dr Lee Romer College of Health and Life Sciences Brunel University UB8 3PH Uxbridge UNITED KINGDOM lee.romer@brunel.ac.uk Dr Samuel Verges HP2 Laboratory UF Recherche Exercice Hop Sud Avenue Kimberley Echirolles CEDEX 38434 Grenoble FRANCE sverges@chu-grenoble.fr

202

Answers to evaluation questions

Please find all correct answers in bold below

Lung Mechanics: Multi-scale Modeling - Dr Jason H.T. Bates

1. Inverse models of lung mechanics necessarily must simple (i.e. have few adjustable parameters)

because:

a. A given data set typically will not allow many parameters to be evaluated within narrow

limits.

b. Having too many different parameters can make it difficult to understand their various

physiological interpretations.

c. There are invariably many different plausible ways that the numerous compartments of a

complicated model can be arranged, and there is no way of knowing from the data which one

is the best.

d. All of the above.

2. Resonant frequency occurs when the imaginary part of impedance (the reactance) is:

a. Unity

b. Zero

c. Infinity

d. Negative

3. The following are potential mechanisms underlying airway hyperresponsiveness:

a. Increased mass of airway smooth muscle

b. Thickening of the airway epithelium

c. Increased access of an inhaled agonist to the airway smooth muscle

d. All of the above

4. Complicated forward models of the lung (i.e. simulation models) are most useful for:

a. proving hypothesis

b. fitting to experimental data

c. disproving hypotheses

d. none of the above

Measuring respiratory mechanics: invasive and noninvasive systems

- Prof. Andrea Aliverti

1. Dynamics refers to

a. The relationships between airflow and volume

b. The relationships between pressure and flow

c. The relationships between volume and pressure

d. None of the above

2. The forced oscillation technique measures

a. The action of the respiratory muscles

b. The impedance of the respiratory system

c. The pressure – flow relationship of the respiratory system at different lung volumes

d. None of the above

3. When integrating airflow to obtain volume

a. Drift is present and is always positive

b. Drift is present and is always negative

c. Drift can be either positive or negative

d. Drift is always negligible

4. Chest wall volume can be measured by

a. Opto-electronic plethysmography

b. Integration of airflow measured by a pneumotachometer

c. Respiratory magnetometers

d. None of the above

5. Volumetric flow rate is

a. equal to flow velocity

b. equal to density / time

c. equal to flow velocity x area of the tube

d. measured by an hot-wire anemometer

Using respiratory mechanics to explain and interpret spirometry – Prof. John Gibson

1. During forced expiration:

a. lung recoil pressure exceeds alveolar pressure

b. intraairway pressure at points of flow limitation exceeds alveolar pressure

c. according to the Bernoulli theorem, flow velocity decreases as air passes from a wider to a

narrower airway

d. equal pressure points (EPPs) in the airways may differ from the sites of maximum

dynamic narrowing

e. audible wheezing implies abnormal airway narrowing.

2. The isovolume pressure-flow (IVPF) curve:

a. can be obtained during a single forced expiratory manoeuvre

b. has a slope equivalent to airway resistance

c. in normal subjects shows effort-dependence at low lung volumes eg 75% FVC expired

d. shows the minimum alveolar pressure needed to generate maximum flow at a given lung

volume

e. in COPD shows a plateau of flow on inspiration

3. In patients with COPD:

a. Maximum flow volume curves show greater reduction of inspiratory than expiratory flow

b. thoracic gas compression during forced expiration has a greater effect on measurements

of FEV1 than in normal subjects.

c. Maximum flow-volume curves help to distinguish the condition from asthma

d. concavity of the descending limb of the maximum flow-volume curve (ie convexity towards

the volume axis) is diagnostic of airway obstruction.

e. lung recoil pressure is usually increased

4. In patients with extrathoracic airway obstruction:

a. maximum expiratory flow is usually reduced more than maximum inspiratory flow

b. a plateau of maximum expiratory flow towards the end of forced expiration is characteristic

c. intrathoracic flow limitation during forced expiration does not occur

d. peak expiratory flow is typically reduced by a greater proportion than FEV1

e. the intensity of stridor indicates the severity of airway narrowing