Systemic cardiovascular effects of volatile and intravenous anesthetics: evaluation in the time domain, the frequency domain and the pressure-volume plane.

Université Libre de Bruxelles Faculté de Médecine

Laboratoire de Physiologie

Directeur : Professeur Robert Naeije

Yvon L.J.M. Deryck

Systemic cardiovascular effects of volatile and intravenous anesthetics:

evaluation in the time domain, the frequency domain

and the pressure-volume plane.

X ī n :

Chinese character for heart.

The Chinese heart is tradionally regarded as the organ of thinking and reasoning, as well as feeling.

Front cover

Dit proefschrift is opgedragen aan : Ce travail de thèse est dédié à : This thesis is dedicated to :

Louis - “Lodewijk” - Deryck

June 12th, 1930 – May 1st, 2012

Dear Timothy and Katya,

everyone needs to realize that :

“The history of mankind is an immense sea of errors in which a few obscure truths may here and there be found.”

Cesare de Beccaria (1738-1794),

Italian jurist, philosopher and politician.

Systemic cardiovascular effects of volatile and intravenous anesthetics : evaluation in the time domain, the frequency domain

and the pressure-volume plane.

Author : Yvon L.J.M. Deryck Promotor : Prof. Dr. R. Naeije

I. Acknowledgements.

II. Summary.

III. Résumé.

IV. Abbreviations.

V. Chapter 1 : Introduction.

VI. Chapter 2 : Evaluation of cardiovascular performance.

VII. Chapter 3 : Systemic vascular effects of isoflurane versus propofol anesthesia in dogs.

VIII. Chapter 4 : Differences in left ventricular-arterial coupling during sevoflurane and propofol anesthesia in dogs.

IX. Chapter 5 : General discussion and conclusions, including clinical implications.

X. Coda : A couple of personal notes on my experiences and perspectives concerning anesthesia related cardiovascular research.

XI. Bibliography.

XII. Additional proposition (« Thèse annexe »).

XIII. Membres du Jury de thèse.

XIV. Curriculum vitae.

I . A c k o w l e d g e m e n t s

This work could only reach an endpoint thanks to the contributions, in one way or another, of many people.

In first instance, my parents Louis Deryck and Mariette Lacus, who gave me the opportunity to start with and to conclude a long academic education. Access to and the benefit of high level education is the most valuable heritage one can get from his parents.

Robert Naeije suggested me to perform research in his laboratory, a suggestion which I accepted with inspiration and great aspirations. I owe him my respect for his endless engagement and patience during the slow progress of this work.

Serge Bimioulle has played a crucial role during the entire process of this work. He developed the software for data acquistion and impedance calculation. He also took an essential part in guiding the experiments, the data analysis, the writing and submission of the papers. He implemented the “single beat estimation” of the end-systolic elastance and as such he is the mentor of this work on left ventricular – arterial coupling.

Marco Maggiorini was my co-worker during the first series of experiments, while Didier De Cannière performed the surgical instrumentation. Luc De Baerdemaeker and Kristine Fonck were my fellow workers during the second series of experiments.

Michel Baurain and Pierre Nokerman, my colleagues at the “Hors Quartier”

Anesthesiology Department of the Clinique Universitaires de Bruxelles - Hôpital Erasme, boosted my perseverance and were supportive in many aspects. Just to name a few: they improved the syntax of my drafts in the language of Molière and they took care of my clinical workload, thereby creating non-protected research time.

The following people were of great help, either for their scientific advice or moral support: Rudy Colson, Patricia Ewalenko, Pierre Pandin, Anita Verleije, Claude Sadis, Patrick Segers, Arlette Vandesteene and Jing Zhao.

Grants for material support were provided by the Belgian Foundation for Cardiac Surgery and by the Fonds de la Recherche Scientifique Médicale (grants 3.4537.91 and 3.4551.05).

Due to some inappropriate and misplaced ambition, I accepted an appointment as an expert witness in a court case. Because of my misconceptions concerning human nature in trial cases, this appointment caused a lot of distraction from my scientific endeavours. Fortunately Jan Van Walleghem, surgeon and LL.B., introduced me to the

‘art and science’ of medical court trials and he was always available for advice.

During more than a decade my life ran into a professional and personal roller-coaster.

Vital to me were the unexpected phone calls, text messages, emails or post cards from nearby and distant friends : Trees Aerts, Gilbert Bejjani, Le Hong Chinh, Jean-Marie Mathues, Evert Rulf, Laszlo Szegedi, Dries van der Woerd and Ivo Van Puyenbroeck.

Ma en Pa^† ne gruute merçi

Robert, Serge, Marco, Didier, Luc and Ariza,

Kristine, Michel,

Pierre, Rudy, Patricia,

Pierre, Anita, Claude, Patrick, Arlette, Jing,

Jan, Trees, Gilbert,

Chinh, Jean-Marie,

Evert, Laszlo,

Dries, Ivo :

Xièxiè.

I I . S u m m a r y

Cardiovascular stability is of prime importance in order to maintain homeostasis during anesthesia and intensive care, and to reduce cardiovascular perioperative morbidity and mortality.

General anesthesia does have profound cardiovascular effects, and the end result is usually a decrease in arterial pressure, with the potential of inadequate organ perfusion and consequently organ damage. Therefore, elucidation of the mechanisms of

cardiovascular effects of general anesthesia is important in order to prevent and/or to treat adequately the cardiovascular perturbations, and to perform the optimal choice of the anesthetic management. Anesthetic management for the patient presenting with cardiovascular alterations relates essentially to the question of a volatile anesthetic based regimen versus a propofol based anesthetic regimen.

A traditional hemodynamic investigation includes the measurement of heart rate, systemic and pulmonary arterial pressure, the filling pressures of the heart and cardiac output. These measurements allows for the calculation of systemic vascular resistance in order to evaluate arterial tone. However, calculated systemic vascular resistance cannot discriminate between passive (flow-dependent) and active (tone-dependent) changes in arterial pressure. Changes in arterial tone must be assessed by constructing pressure-flow plots.

Neither calculated systemic vascular resistance nor pressure-flow plots takes into account the pulsatile nature of the circulation. In order to do so, one has to measure instantaneous pressure and flow waves, perform harmonic analysis on both waves and calculate vascular impedance spectra.

The cardiovascular system is a mechanical system in which two components are functionally coupled: there is an energy transfer between the energy source, i.e. the left ventricle, and its mechanical load, i.e. the arterial tree. An alteration in one of these components necessitates an appropriate alteration in the other component in order to maintain optimal coupling, i.e. maximal energy transfer between the two elements. In the pressure-volume plane the left ventricle and the arterial tree are considered to be two elastic chambers in series. The performance of the left ventricle is quantified by the end-systolic elastance, while the load of the arterial tree is quantified by the effective arterial elastance. The ratio of end-systolic elastance to effective arterial elastance relates ventricular-arterial coupling to either maximisation of stroke work or either to maximisation of mechanical efficiency (i.e. the ratio of mechanical power output to cardiac oxygen consumption).

In the first experiment we investigated the systemic vascular effects of isoflurane versus propofol anesthesia in dogs using a traditional hemodymamic approach, measurement of instantaneous aortic flow and pressure with subsequent calculation of aortic input impedance spectra, and construction of pressure-flow plots generated by gradual reduction of venous return. Calculated systemic vascular resistance could not detect differences in arteriolar tone between isoflurane and propofol, whereas pressure-flow plots did: compared with isoflurane, propofol better maintained aortic pressure at all levels of flow, except at the lowest level of flow. Impedance spectra demonstrated a decreased pulsatile load and less energy losses in pulsations with propofol as compared with isoflurane.

In the second experiment we investigated the effects of escalating doses of sevoflurane and propofol anesthesia on arterial mechanical properties and left ventricular-arterial coupling in the dog. Arterial mechanics were assessed by traditional hemodynamics, aortic input impedance spectra, and pressure-flow plots generated by rapid caval inflow reduction. Left ventricular-arterial coupling was assessed as the ratio of end-systolic elastance to effective arterial elastance. The end-systolic elastance and the effective

arterial elastance were obtained from left ventricular pressure and aortic flow data using a ‘single-beat’ estimation method. Traditional hemodynamics and pressure-flow plots demonstrated that sevoflurane causes a limited arteriolar vasodilation and causes arterial hypotension essentially by a decrease of cardiac output. Propofol insignificantly

decreases cardiac output, but is an “actual” arteriolar dilator. The impedance spectra demonstrated that sevoflurane and propofol do have different effects on the elastic properties of large conduit arteries. Sevoflurane increased the characteristic impedance and reduced arterial compliance, indicating an increased physical elastance of the arterial tree. Propofol caused an insignificant increase of the characteristic impedance and the arterial compliance remained unaltered, suggesting that propofol does have a beneficial effect on the elastic properties of the arterial tree, thereby confirming the conclusion of the first experiment (i.e. a decreased pulsatile load with propofol).

Sevoflurane impaired ventricular-arterial coupling by decreasing end-systolic elastance and increasing effective arterial elastance. Propofol maintained left ventricular-arterial coupling: the end-systolic elastance and effective arterial elastance remained unchanged and as consequence the ratio of end-systolic elastance to effective arterial elastance did not change. All results taken together we conclude that sevoflurane decreases cardiac output and left ventricular contractility, and increases the pulsatile and total load to the left ventricle. Propofol maintains cardiac output and left ventricular contractility, induces an arterial dilatation but without affecting the pulsatile and total load to the left ventricle.

These results, obtained in dogs, suggest that propofol, compared to volatile anesthetics, is an anesthetic, which can better preserve hemodynamic stability and homeostasis in the cardiovascular compromized patient undergoing surgery.

I I I . R é s u m é

Evaluation des effets cardiovasculaires systémiques des agents anesthésiques dans le domaine du temps, le domaine de la fréquence et le plan pression-volume.

La stabilité cardiovasculaire est d’une importance prioritaire pour maintenir

l’homéostasie pendant l’anesthésie et le séjour aux soins intensifs, et pour réduire la morbidité et mortalité cardiovasculaire pendant la période péri-opératoire.

L’anesthésie générale exerce des effets marqués sur le système cardiovasculaire.

Généralement une hypotension artérielle systémique est observée, avec la possibilité d’une hypoperfusion des organes vitaux et ultérieurement des lésions de ces mêmes organes. Donc l’éclaircissement des mécanismes des effets cardiovasculaires de l’anesthésie générale est important pour prévenir et traiter les perturbations cardiovasculaires, et pour effectuer le choix optimal de la gestion anesthésique.

La question de la gestion anesthésique chez le patient présentant une fonction

cardiovasculaire altérée se traduit essentiellement par le choix de l’anesthésie soit basée sur un agent volatile soit basée sur le propofol intraveineux.

Une exploration traditionnelle de l’hémodynamique comprend le mesure de la

fréquence cardiaque, des pressions artérielles systémique et pulmonaire, des pressions de remplissage et du débit cardiaque. Ces mesures permettent de calculer la résistance vasculaire systémique de manière à évaluer le tonus artériel. Cela dit, la résistance vasculaire systémique calculée ne peut pas faire la différence entre des changements actifs (changements du tonus artériel) ou passifs (changements des débits) de la pression artérielle systémique. Les changements du tonus artériel doivent être évalués par des courbes pression - débit.

Ni les résistances vasculaires systémiques ni les courbes pression débit ne tiennent compte de la nature pulsatile de la circulation. L’exploration des effets pulsatiles requiert tout d’abord la mesure des pressions instantanées et des débits instantanés.

En seconde lieu, ces signaux doivent subir une décomposition harmonique (analyse de Fourier), pour afin de pouvoir calculer le spectre d’impédance vasculaire.

Le ventricule gauche et le système artériel sont deux éléments d’un système mécanique, dans lesquels il y a un transfert d’énergie entre la source d’énergie et sa charge.

Une modification dans un des éléments nécessite une modification appropriée dans l’autre élément pour maintenir un couplage optimal entre les deux éléments, c'est-à-dire un transfert maximal d’énergie. Dans le plan pression volume, le ventricule gauche et l’arbre artériel sont considérés comme deux chambres élastiques en série.

La performance du ventricule gauche est quantifiée par l’élastance ventriculaire télésystolique, et la charge du système artériel est quantifiée par l’élastance artérielle effective. Le rapport entre l’élastance ventriculaire télésystolique et l’élastance artérielle effective permet de situer le « couplage ventriculo-artériel » soit en termes de

maximisation du travail ventriculaire ou soit en termes d’ efficience mécanique.

L’efficience myocardique est définie comme un rapport entre la puissance ventriculaire produite et l’oxygène consommé.

Dans la première expérimentation, nous avons étudié les effets vasculaires sur la circulation systémique du chien d’une anesthésie inhalatoire à l’isoflurane versus une anesthésie au propofol, ceci au moyen d’une exploration hémodynamique traditionnelle, les spectres d’impédance aortique et les courbes pression débit étant générées par une réduction graduelle du retour veineux. Les résistances vasculaires systémiques calculées n’ont pas décelé de différences de tonus artériolaire entre les effets d’une anesthésie inhalatoire à l’isoflurane et les effets d’une anesthésie intraveineuse au propofol. Par contre les courbes pression-débit démontrent une différence : comparé à

l’anesthésie à l’isoflurane, l’anesthésie au propofol maintientt mieux la pression aortique à tous les niveaux de débit sanguin sauf aux débits les plus bas. Les spectres d’impédance démontrent une charge pulsatile réduite et des pertes d’énergie réduites avec le propofol par rapport à l’isoflurane.

Dans la seconde expérimentation chez le chien, nous avons étudié les effets de doses croissantes de deux agents anesthésiques généraux, le sevoflurane et le propofol, sur les caractéristiques mécaniques du système artériel et le couplage ventriculo-

artériel systémique. La mécanique artérielle était étudiée par une exploration hémodynamique traditionnelle, les spectres d’impédance aortique et les courbes pression-débit étant générées par une réduction rapide du retour veineux. Le couplage ventriculo-artériel systémique était calculé par le rapport entre l’élastance ventriculaire télésystolique et l’élastance artérielle effective. L’élastance ventriculaire télésystolique et l’élastance artérielle effective ont été estimées à partir de la pression ventriculaire gauche et du débit aortique instantané en appliquant une méthode dite de

« single beat ». L’hémodynamique traditionnelle et les courbes pression - débit

démontrent que le sevoflurane provoque une vasodilatation artériolaire limitée et que la cause principale de l’hypotension artérielle est une réduction du débit cardiaque.

Le propofol réduit le débit cardiaque d’une manière non significative, mais est un vasodilatateur artériolaire réel. Les spectres d’impédance montrent que le sevoflurane et le propofol ont des effets différents sur les caractéristiques élastiques des grosses artères à conduction. Le sevoflurane augmente l’impédance caractéristique et réduit la compliance artérielle, indiquant une augmentation de l’élastance physique de l’arbre artériel. Le propofol provoque une augmentation non significative de l’impédance caractéristique, mais la compliance artérielle reste inchangée. Ces résultats suggèrent que le propofol aurait un effet favorable sur les propriétés élastiques de l’arbre artériel, et donc confirment les conclusions de la première expérimentation, c’est-à-dire une

charge pulsatile réduite avec le propofol. Le sevoflurane dégrade le couplage

ventriculo-artériel à la suite d’une réduction de l’élastance ventriculaire télésystolique et d’une augmentation de l’élastance artérielle effective. Le propofol maintient le

couplage ventriculo-artériel. L’élastance ventriculaire télésystolique et l’élastance artérielle effective restent par contre inchangées. Par conséquent, le rapport entre les deux élastances ne change pas. Sur base de ces résultats, nous concluons que le sevoflurane réduit le débit cardiaque et la contractilité du ventricule gauche, et

augmente la charge pulsatile et totale sur le ventricule gauche. Le propofol maintient le débit cardiaque et la contractilité du ventricule gauche, et induit une dilatation artérielle sans altérer la charge pulsatile et totale sur le ventricule gauche.

Ces résultats, obtenus chez le chien, suggèrent que le propofol, comparé aux

anesthésiques volatiles, est un anesthésique qui permet de mieux préserver la stabilité hémodynamique et l’homéostasie chez le patient présentant une fonction

cardiovasculaire restreinte et devant bénéficier d’un acte chirurgical.

I V . A b b r e v i a t i o n s AoQ : aortic flow

!

AoQ : mean aortic flow AoP : aortic pressure

!

AoP, : mean aortic pressure c0 : pulse wave velocity Ca : total arterial compliance CME : Cardiac Mechanical Efficiency CO : Cardiac Output

CVP : Central Venous Pressure Ea : Effective Arterial Elastance Ees : End-systolic Elastance Emax : maximum elastance E(t) EF : Ejection Fraction

EDV : End-Diastolic Volume ESV : End-Systolic Volume

ESPVR : End-Systolic Pressure-Volume Relationship E(t) : time-varying elastance

EW : External (Mechanical) Work

f : frequency

fmax : frequency at first local maximum of SVZ modulus (: Zmax) fmin : frequency at first local minimum of SVZ modulus (: Zmin)

g : gram

HR : Heart Rate

IV : Intravenous

kg : kilogram

!

l : length of a tube or vessel segment

L : Liter

LAP : Left Atrial Pressure

LVAC : Left Ventricular- Arterial Coupling LVEDP : Left Ventricular End-diastolic Pressure LVEDV : Left Ventricular End-diastolic Volume MAC : Minimum Alveolar Concentration MAP : Mean Arterial Pressure

min : minute

mmHg : millimeters of mercury

N : the total number of harmonics computed PaCO2 : arterial carbon occluded

PaO2 : arterial oxygen tension

PADP : Pulmonary Artery Diastolic Pressure PAOP : Pulmonary Artery Occlusion Pressure Pao(t) : instantaneous aortic pressure (at time t) Pdias : diastolic arterial blood pressure

PE : Potential Energy

Pee : venricular pressure at end ejection Pes : ventricular pressure at end systole Ped : ventricular pressure at end diastole pHa : arterial pH

Psys : systolic arterial blood pressure

Pin, Pout : pressure at the inflow and outflow of a tube or vessel segment Pmax : peak systolic pressure of an isovolumic beat

PP : pulse pressure (= P – P )

PTF : Pressure Transfer Function

P(t) : instantaneous (left) intraventricular pressure PRSW : Preload Recruitable Stroke Work

PvO2 : mixed venous oxygen tension P-V : Pressure-Volume

PVA : Pressure Volume Area

PVR : Pulmonary Vascular Resistance

Pv(t) : instantaneous (left) intraventricular pressure PWI : Pressure Work Index

PWV : Pulse Wave Velocity RAP : Right Atrial Pressure

Q : volume flow

Qao(t) : instantaneous aortic flow (at time t) Qn : the amplitude of the n-th flow harmonic r : radius of a tube or vessel sefment R : hydraulic or vascular resistance Rc : Reflection coefficient

Rin : input resistance SV : Stroke Volume SD : Standard Deviation

SE, SEM : Standard Error of the Mean SVR : Systemic Vascular Resistance SVZ : Systemic Vascular Input Impedance

SW : Stroke Work

TIVA : Total Intravenous Anesthesia

T : heart period

Tb : time of beginning of the cardiac cycle Te : time of end of cardiac cycle

Tmax : time from the onstet of end systole to reach Emax

Ved : end-diastolic volume Ves : end-systolic volume

Vmax : maximal velocity of muscle fiber shortening

Vo2 : left ventricular (myocardial) oxygen consumption (per beat)

!

W ˙ _osc : oscillatory (external) flow (or hydraulic) power

!

W ˙ _st : steady (external) flow (or hydraulic) power

!

W ˙ _tot : total (exernal) hydraulic power

V(t) : instantaneous (left) intraventricular volume Z0 : input resistance

Zc : characteristic impedance

Zmax : SVZ modulus at first local maximum Zmin : SVZ modulus at first local minimum

Zn : the modulus of the n-th harmonic of the aortic input impedance Z(ω) : Input Impedance at angular frequency ω

Γ : reflection coefficient (wave -) η : (dynamic) viscosity of a fluid

ρ : blood density

θn : the phase angle of the n-th harmonic of the aortic input impedance

V . C h a p t e r 1 Introduction.

In human clinical practice, induction and maintenance of general anesthesia is usually accomplished either by inhalation of a volatile anesthetic or by an intravenous injection of an intravenous anesthetic. Both inhaled and intravenous anesthetics can produce the desired clinical endpoints for surgery, i.e. unconsciousness and absence of movement to noxious stimulation [67, 127, 195]. Since the introduction of barbiturates, intravenous induction of general anesthesia has gained widespread acceptance [96]. However, notwithstanding the introduction of short acting intravenous anesthetics, propofol for instance, and the introduction of the concept of total intravenous anesthesia (TIVA), maintenance of general anesthesia with volatile anesthetics is still very popular among most anesthesiologists [161].

General anesthetic drugs do have profound cardiovascular effects, and general anesthesia is commonly associated with a decrease in arterial blood pressure [105].

According to the hydraulic equivalent of Ohm’s Law the reduction in pressure is caused either by a reduction in flow, i.e. cardiac output, or a reduction in arterial tonus or a combination of both. The reduction in cardiac output is the consequence of either a reduced intrinsic myocardial contractility or a reduced filling of the heart, or a combination of both.

Potential mechanisms of this myocardial depression and/or vasodilation, acting in isolation or in combination, includes [62, 206, 234, 257] :

- a direct effect upon the effector organ, i.e. the cardiac myocytes and vascular smooth cells,

- an indirect effect through the autonomic nervous system and neurohumoral control mechanisms,

- an effect upon the endothelium or endocardium.

The central nervous system is the “target organ” for general anesthesia. However modulation of the autonomic nervous system, decreased orthosympathetic or increased parasympathetic outflow, always contributes to a depressed cardiovascular state.

Volatile and intravenous anesthetics differs as to the extend of orthosympathetic depression, ketamine being the sole IV anesthetic exhibiting orthosympathetic excitation [96] .

Volatile and intravenous anesthetics acts directly on the cardiac myocytes through interference with cellular calcium handling. Volatile anesthetics depress contractile performance through a decreased intracellular calcium availability and a decreased myofibrillar calcium sensitivity [13, 106, 107]. Propofol negative inotropic effect at high concentrations results solely from a decreased intracellular calcium availability [47].

Volatile anesthetics demonstrate a dual effect on vascular smooth muscle cells in vitro:

relaxation and constriction. These effects are the result of either a direct effect on the vascular smooth cell or an indirect effect (usually by inhibition of vasorelaxation) [2, 162, 235-237, 256]. The direct relaxing effects are due to a decreased intracellular calcium concentration and a decreased calcium sensitivity [2, 112, 257]. Propofol vasodilating action involves both a calcium blocking effect and calcium desensitizing action [257] .

Volatile and intravenous anesthetics also have an effect upon the vascular endothelium:

as a consequence general anesthetic agents also modulates vascular smooth muscle tone through actions on the endothelium [2, 257]. The endocardial endothelium, a continuum with the vascular endothelium, constitutes a structure that actively regulates ventricular performance [54, 55]. However, modulation of ventricular function by general

anesthetic agents through interference with the endocardium has only been reported for thiopental [16] .

Cardiovascular diseases represent a world wide major medical problem ^#. As a consequence the number of patients with cardiovascular impairment presenting for major surgery is increasing. Moreover, some disease states, e.g. diabetes, also affect cardiovascular performance through their impact upon autonomic (orthosympathetic) nervous system or/and endothelial function [84, 257]. Therefore, investigation of the cardiovascular effects of general anesthetics is of prime importance in order to reduce cardiovascular perioperative morbidity and mortality.

In addition, the pathophysiological approach to cardiovascular disease has been changed over the years. Cardiovascular diseases are now being increasingly recognized as ventricular-arterial coupling disease, rather than a disease of one particular aspect of the cardiovascular system, e.g. failure of the left ventricle or the arteriolar vasoconstriction [20, 36, 38, 115, 116]. As a consequence, adequate methods should be developed and applied to investigate the cardiovascular effects of general anesthetics, which are not revealed by a traditional approach.

A traditional hemodynamic investigation, both in the clinical and laboratory setting, includes the measurement of heart rate, systemic and pulmonary arterial pressure, the filling pressures of the heart and cardiac output. The measurement of the latter three variables is accomplished by the insertion of a flow-directed balloon-tipped catheter (the so-called “Swan-Ganz catheter” or pulmonary artery catheter) via a central vein and positioning into a branch of the pulmonary artery [248]. This catheter allows for the measurement of cardiac output by the thermodilution method [73, 79]. In this approach pressure measurements are accomplished by the use of fluid filled catheters,

transmitting the pressure signal to a pressure transducer, which converts the mechanical signal into an electrical signal. These fluid filled catheters do have poor system dynamic characteristics, which allows only for measurement of “mean pressures” with

reasonable accuracy [43, 169, 196].

#http://www.who.int/cardiovascular_diseases/en/

http://www.who.int/mediacentre/factsheets/fs317/en/index.html

Simultaneous measurement of the systemic and pulmonary arterial pressure and the cardiac output allows for the calculation of systemic (SVR) and pulmonary (PVR) vascular resistance. Although SVR and PVR are traditionally used for evaluating arterial changes, calculating vascular resistance does have limitations. First, calculated vascular resistance cannot discriminate between passive (flow-dependent) and active (tone-dependent) changes in arterial pressure [137]. Changes in arterial tone are better assessed by comparing pressures at several levels of flow. Second, because of the pulsatile nature of the circulation – the heart is not a steady pump and the blood vessels are not rigid tubes – complex pressure and flow waves are generated in the arterial bed.

These complex waves cannot be adequately described by one single number, i.e. the

“mean”. Characterization of these complex waves requires the measurement and recording of the instantaneous pressure and flow waves with sufficiently high accuracy and subsequent mathematical analysis, namely harmonic analysis (Fourier

decomposition) [155, 169]. As a consequence calculated vascular resistance offers only a partial description of the forces opposing blood flow across a vascular bed.

Cardiac output is a global indicator of ventricular performance, but is poor indicator of intrinsic ventricular contractility, because cardiac output is also dependent on heart rate and loading conditions of the heart (preload and afterload) [21, 22]. The concept of the end-systolic pressure-volume relationship offers the opportunity to evaluate myocardial contractility relatively independent of loading conditions of the heart [214, 215]. In this approach myocardial contractility is characterized by the end-systolic elastance Ees, which is the slope of the line connecting the end-systolic points in pressure-volume loops generated during different level of preload or afterload. This methodology necessitates the simultaneous measurement of ventricular instantaneous pressure and volume at different level of preload or afterload. The measurement of ventricular volume is technically difficult and requires expensive technology. Therefore a “single beat” method has been developed in order to estimate Ees [27, 247, 249, 253] .

The left ventricle and the arterial tree are two elements of a mechanical system, in which there is an energy transfer between the energy source and its mechanical load [169]. An alteration in one the elements necessitates an appropriate alteration in the other element in order to maintain optimal coupling, i.e. maximal energy transfer between the two elements. In the pressure-volume plane the left ventricle and the arterial tree are considered to be two elastic chambers in series. The performance of the left ventricle, i.e. the energy source, is quantified by Ees. The load presented by the arterial tree is quantified by the effective arterial elastance Ea : the slope of the line between the end-diastolic point and the end-systolic point in a pressure-volume loop [125, 244]. The ratio of Ees to Ea creates the opportunity to describe the ventricular- arterial coupling in terms of mechanical efficiency, defined as the ratio of mechanical power output and myocardial oxygen consumption [32].

The objective of this thesis is to compare the cardiovascular effects of volatile anesthetics versus an intravenous anesthetic, namely propofol, in the time-domain (steady flow), the frequency domain (pulsatile flow) and the pressure-volume (PV) plane. We hypothesized that:

1. differences in cardiovascular effects are better detected using pressure- flow plots, frequency domain techniques and analysis in the PV plane,

2. volatile anesthetics and propofol would impair left ventricular – arterial coupling by different mechanisms, but that propofol better maintains this left ventricular-arterial coupling.

Chapter 2 of this thesis treats the evaluation of the systemic circulation. First, the determinants of (left) ventricular pump function are briefly described. Second, the assessment of preload is commented. In a third section, methods for the characterization of the arterial load are presented and discussed, including time-domain and

frequency-domain methods. Subsequently, methods for the assessment of myocardial contractility are commented: load-dependent indices are mentioned, but the focus is on the end-systolic pressure-volume relationship. This PV analysis also introduces another index to describe pump performance, namely ventricular work. In the final section, ventricular – arterial interaction is analyzed in the PV plane.

Chapter 3 of this thesis reports our results on the vascular effects of isoflurane versus propofol anesthesia in dogs.

Chapter 4 of this thesis reports on the differences in left ventricular-arterial coupling during sevoflurane and propofol anesthesia in dogs.

Chapter 5 of this thesis comments and concludes on the data reported in chapter 3 and chapter 4.

The data reported in Chapter 3 and 4 are based on the following papers:

Deryck Y, Brimioulle S, Maggiorini M, De Canniere D, Naeije R : Systemic vascular effects of isoflurane versus propofol anesthesia in dogs, Anesthesia &

Analgesia, 1996, 83: 958-964,

Deryck Y, Fonck K, De Baerdemaeker L, Naeije R, Brimioulle S : Differential effects of sevoflurane and propofol anesthesia on left ventricular-arterial coupling in dogs, Acta Anesthesiologica Scandinavica, 2010, 54 : 979-986.

Further mathematical analysis of the data related to the first paper resulted in the following paper:

Segers P, Verdonck P, Deryck Y, Brimioulle S, Naeije R : Pulse pressure method and the area method for the estimation of total arterial compliance in dogs : sensitivity to wave reflection intensity, Annals of Biomedical

Engineering, 1999, 27 : 480-485.

The author of this thesis also co-authored the following paper, concerning the cardiac effects of propofol in a clinical setting :

Schmidt C, Roosens C, Struys M, Deryck Y, Van Nooten G, Colardyn F, Van Aken H, Poelaert J : Contractility in humans after coronary artery surgery, Echocardiographic assessment with preload-adjusted maximal power, Anesthesiology, 1999, 91 : 58-70.

V I . C h a p t e r 2

Evaluation of cardiovascular performance.

2.1. Determinants of ventricular pump performance.

2.2. Assessment of preload.

2.3. Assessment of arterial load.

2.3.1. Time domain techniques.

2.3.1.1. Calculated (systemic) vascular resistance.

2.3.1.2. Arterial pressure-flow plots.

2.3.2. Frequency domain techniques.

2.3.2.1. Arterial input impedance spectra.

2.3.2.1.1. Quantification of wave reflections.

2.3.2.1.2. Practical aspects concerning vascular impedance determinations.

2.3.2.2. Pressure transfer function.

2.3.3. Effective arterial elastance.

2.4. Assessment of (intrinsic) contractility.

2.4.1. Load-dependent indices of contractility.

2.4.2. The end-systolic pressure-volume relationship.

2.4.2.1. The time-varying elastance model of the ventricle.

2.4.2.2. Practical aspects concerning the determination of Ees

2.4.3. Preload recruitable stroke work.

2.5. Ventricular – arterial interaction and mechanoenergetic efficiency.

2.5.1. Ventricular work and power.

2.5.2. Cardiac energetics and efficiency.

2.5.3. Ventricular-arterial coupling and effective arterial elastance.

2.5.4. Computational aspects for determining ventricular-arterial.

coupling and cardiac mechanical efficiency.

The cardiovascular system (or circulation) is a transportation system that supplies all the cells of the body, delivering essential materials (oxygen, nutrients, signaling molecules) and carrying away the waste products of metabolism [154]. In order to do so the system is composed of a pump (the heart), some circuitry to distribute (the arteries) and collect (the veins) the blood to and from the peripheral tissues, and a vast capillary network to allow exchange between the blood and the tissues. The heart is a dual pump that drives the blood in two serial circuits: the systemic and pulmonary circulations.

This thesis is concerned with the effects of general anesthetics upon the systemic circulation, but the principles outlined here are equally applicable to the pulmonary circulation, although one must take into account some essential differences between the two circulations [154] :

- the geometry of the left and the right ventricle,

- the structural differences between the pulmonary and systemic arteries, - the pulmonary circulation is a “low pressure” circuit (mean pressure in the

pulmonary artery approximates 15 mmHg versus 95 mmHg in the aorta).

The foregoing entails a dual approach for the evaluation of cardiovascular function.

Biochemical measurements relates to adequate cellular functioning in peripheral tissues.

Physical measurements relates to the transportation function itself, i.e. “material flux”

[53], and to the structural characteristics of the system. Measuring material flux is usually accomplished through measurement of flow; flow being the displacement of volume per unit time. Depending on the clinical or experimental setting the chemical or the physical approach will be emphasized, but both approaches are complimentary to each other.

The function of the heart as a pump is to eject a certain amount of blood into the arterial tree in order to provide adequate organ perfusion. The pump performance is quantified by its output: the cardiac output.

!

CO=SV"HR"k (Eq.2.1)

with: CO : cardiac output (l/min) SV : stroke volume (ml/beat) HR : heart rate (beats/min) k : conversion factor (=10^-3)

Given equation 2.1, one will consider ventricular pump performance as the ability to generate stroke volume. Although equation 2.1 is mathematically correct, experimental investigations has demonstrated that increases in heart rate either produces only a limited increase cardiac output, or are detrimental at high heart rates, essentially because heart rate and stroke volume are not independent variables [139].

2.1. Determinants of ventricular performance.

The determinants of ventricular pump performance are considered as preload, afterload and contractility [258]. The concepts of preload, afterload, and (intrinsic) contractility are derived from studies on isolated (mammalian papillary) muscle preparations and as such are difficult to apply straightforward to a three-dimensional ventricle [15, 120, 139, 154, 155, 169, 214]. Nonetheless, because these concepts are part of clinicians daily vocabulary and are often defined differently [177], one needs an “operational definition” of these concepts [120, 214].

Contractility literally means ability of shortening or developing increased force [154, 228]. This definition implicitly acknowledges two types of contraction : isotonic and isometric contractions. According to this definition quantification of contractility should be straightforward: measuring shortening during an isotonic contraction or measuring force during an isometric contraction. However, experiments indicate that contractility is dependent on the loading conditions of the muscle [111]. Therefore, a (quantitative) definition of contractility requires in first instance a clarification of the terms preload and afterload. The difference between a preload and an afterload essentially depends on when the muscle first interacts with the load (figure 2.1) [120].

Figure 2.1

Preload (PL) and afterload (AL) in a papillary muscle.

A : Resting stage in the intact heart just before opening of the atrioventricular valves.

B : Preload in the intact heart at the end of ventricular filling.

C : Supported preload plus afterload in the intact heart just before opening of the aortic valve.

D : Lifting preload plus afterload in the intact heart : ventricular ejection with decreased ventricular volume.

From : Levy, M. N., Stanton, B. A., & Koeppen, B. M. (2006). Berne & Levy Principles of Physiology (Fourth ed.). Philadelphia: Mosby Elsevier.

Preload is the weight (= force) to which a muscle is subjected before contraction begins, actually it is the passive load that establishes the initial muscle length of the muscle fibers prior to contraction (i.e. an isometric contraction) [120, 228]. Between some limits there is a direct relation between preload and the force generated: the greater the preload the greater the energy generated. This the Frank-Starling mechanism [71, 226].

Preload of the ventricle is considered to be the load on the ventricle at the moment before contraction starts. Left ventricular volume at end-diastole is accepted as the most appropriate measure of ventricular preload [214]. This is the point in the right hand lower corner the pressure-volume loop of figure 2.2, which represent the cardiac cycle in the pressure-volume plane.

Figure 2.2

Diagram of pressure and volume in human left ventricle.

Lower continuous line : resting or diastolic relationship.

Upper continuous line: pressure maxima reached from indicated resting volume under isovolumic conditions.

Dashed line : sequence in typical normal beat.

D : diastole, IC : isovolumic contraction period, EJ : ejection, IR : isovolumic relaxation.

From : Milnor, W. R. (1982). Hemodynamics. Baltimore: Williams & Wilkins

Figure 2.3

The effect of increasing initial length of a cat papillary muscle on the force-velocity relationship, degree of shortening, muscle work and muscle power.

From : Berne, R. M., & Levy, M. N. (1981). Cardiovascular Physiology (Fourth ed.). St. Louis: The C.V. Mosby Company.

Afterload is the force resisting muscle fiber shortening [24, 254]. Practically it is the arrangement of a muscle, in an isolated muscle preparation (figure 2.1), so that in shortening, it lifts a weight from a support, i.e. an isotonic contraction [120, 228]. The extent of shortening and its velocity are inversely related to the afterload (figure 2.3) [71].

Defining afterload as the forces opposing ventricular muscle fiber shortening, implies the forces opposing the ejection of blood from the left ventricle. There is however considerable debate as how to determine these forces, and one can consider two

“schools of thought”. The “cardiac school” refers to afterload as “some ventricular measure of force” that a ventricle must overcome while it contracts during ejection, e.g.

afterload defining as ventricular pressure, implicates that the ventricle plays a part in determining its own afterload, because ventricular pressure is not independent of ventricular contractility [153]. Figure 2.4 also demonstrates that intraventricular pressure varies with time during ejection, so that afterload is not constant.

Figure 2.4

Time function curves of ventricular pressure P_v(t), aortic pressure P_ao(t), atrial pressure P_atr(t), ventricular volume V_v(t), and aortic flow F_ao(t).

From : Sagawa, K., Maughan, L., Suga, H., & Sunagawa, K. (1988). Cardiac Contraction and the Pressure-Volume Relationship. New York: Oxford University Press.

The “vascular school” considers afterload as the “external” factors that oppose ventricular ejection, and this approach regards the arterial system itself as the ventricular afterload [153, 154]. The factors that oppose ventricular ejection are the

viscosity of blood and the viscoelasticity of the arterial system [153, 154]. This thesis will adopt the “vascular definition” of afterload. In doing so, one will adopt the view that the hydraulic load presented to the ventricle is ventricular afterload.

Given the load dependence of muscle shortening or force generation in an isolated muscle experiment, contractility should be defined as a quantitative index, representing the inherent capacity of the muscle to contract independently of preload and afterload.

For an isolated muscle, the maximal velocity of shortening Vmax is a reasonable index of contractility (figure 2.3) [15, 25, 155, 186]. However, given the foregoing discussion concerning afterload and given technical limitations, an unambiguous quantitative definition of contractility may be elusive in both the clinical and experimental setting.

In addition, insisting that contractility is independent of load may be incompatible with the cellular mechanisms of contractility [10, 130, 154, 155, 186].

Contractility is also dependent on the frequency of contraction. Increasing the frequency of contraction progressively increases the force of contraction. This in know as the treppe or staircase phenomenon and was first described by Bowditch for the frog ventricle [15, 155, 186]. As a consequence a contractility index should also take into account heart rate. In the clinical setting and in terms of “macroscopic” cardiovascular mechanics in the intact subject, it nonetheless remains important to separate the effects of a primary increase in load or heart rate from a primary increase in contractility [186].

As for this thesis, one will retain the following working, albeit traditional, definition of contractility: contractility reflects the ability of the myocytes to generate force given a specific load [95] .

Note 1 : The foregoing approach concerning ventricular performance focus on the generation of flow, i.e. the systolic function, and neglects the diastolic function [21, 187, 188, 254]. However one should realize that systolic and diastolic are tightly coupled.

Note 2 : Regarding the systolic properties of the ventricles the terms ‘performance’,

‘function’ and “contractility’ are frequently used interchangeable in the literature. One already pointed out the meaning of the term ‘contractility’. Concerning the other two terms, this thesis will adopt Braunwauld’s terminology [11, 21]:

- Ventricular performance is related to the simple pumping function of the ventricle, as reflected in the cardiac output or cardiac work, expressed per stroke (stroke volume and stroke work) or per minute;

- Ventricular function relates these variables of ventricular performance to some measure of preload.

2.2. Assessment of preload.

As already alluded to, preload represents the volume that produces stretch of the myofibrils and determines sarcomere length prior to contraction [95].

Technical limitations to measuring intracardiac volumes have led reliance upon intracardiac pressures as an indirect indicator of preload [95] [218]. Clinically and experimentally left atrial pressure (LAP), pulmonary artery occlusion pressure (PAOP), pulmonary artery diastolic pressure (PADP), right atrial pressure (RAP), or central venous pressure (CVP) are often used as substitutes for left ventricular end diastolic pressure (LVEDP) and left ventricular end diastolic volume (LVEDV) [205]. Their accuracy in predicting LV preload is determined by the distensibility properties of the ventricle (or ventricular compliance), the integrity of the mitral valve, the presence of normal pulmonary conditions, the integrity of the pulmonic and tricuspid valves, and right ventricular function. The assumption that ventricular distensibility is normal is not a valid assumption in many patient with cardiac disease [205]. In addition the relation between LVEPV and LVEDP is not linear [83, 205, 261].

In a laboratory setting instantaneous left ventricular dimensions can be measured by means of sonomicrometry [88, 214]. Since the introduction of the “conductance catheter” the measurement of left ventricular volume and pressure in the experimental and clinical setting have become possible [7, 9, 214]. Introduction of the “conductance catheter” in the left ventricle also allows the (simultaneous) measurement of left

ventricular volumes and pressure. The applicability of this technique is limited due to its invasiveness, high costs and the requirement of a sophisticated infrastructure.

Another means of assessing left ventricular preload is by application of two dimensional echocardiography and by using appropriate geometric models of the left ventricle one can estimate left ventricular volume [34]. However, recent advances in three

dimensional echocardiography, either transthoracic or transoesophageal, allows for non- invasive accurate measurement of left ventricular volume [131]

Because the investigations reported in this thesis focus on afterload and ventricular- arterial coupling, preload in the reported investigations was assessed by measuring RAP and PAOP.

2.3. Assessment of arterial load.

2.3.1. Time domain techniques.

2.3.1.1. Calculated systemic vascular resistance.

Although the physical laws governing fluid flow were first described by Navier and Stokes, hemodynamic evaluation of cardiovascular function relies heavily on the Poiseuille – Hagen equation [155, 169]. This equation, which is a particular solution of the Navier-Stokes equations, describes the steady, laminar flow of a Newtonian fluid through a cylindrical tube and states that the flow varies directly as the pressure

difference (or gradient) Pin-Pout and the fourth power of the radius of the tube and varies inversely as the length of the tube and the viscosity of the fluid. Poiseuille Law states [155] :

!

Q=(P_in"P_out)#.r⁴

8.$.l (Eq. 2.2)

where Q : volume flow,

Pin – Pout : pressure drop (or : driving pressure) along a tube of length

!

l, r : internal radius of the tube,

η : viscosity of the fluid,

!

l : length of the tube.

The term steady flow signifies the absence of variations of flow in time, i.e. flow is time-invariant. Laminar flow is the type of motion in which the fluid moves as series of individual layers, with each stratum moving at a different velocity from its

neighbouring layers [15, 154]. A Newtonian fluid is one in which the viscosity is constant at all shear rates [87]. Equation 2.2 points out that flow is extremely sensitive to the internal radius of the tube : Q ∝ r⁴.

Rearranging equation 2.2 leads to :

!

P_in "P_out

Q =8.#.l

$.r⁴ (Eq. 2.3)

This equation states that the ratio of driving pressure to flow is a function of the dimension of the tube and the viscosity of the moving fluid. The physical properties of the system determine how large a pressure difference is required to produce a given flow. The ratio of driving pressure to mean flow is thus a measure of the extent to which the system “opposes” or “resists” to flow [155]. In analogy to Ohm law for electricity this ratio is called hydraulic resistance R in fluid mechanics and vascular resistance in cardiovascular physiology [155].

!

R=8.".l

#.r⁴ (Eq. 2.4)

Equation 2.4 can be written as :

!

R= 8

"

#

$ % &

' ( )*) l r⁴

#

$ % &

' ( (Eq. 2.5)

This means that the resistance is determined by the physical properties of the fluid (η) and the geometry of the tube

!

lr⁴

( )

The following table illustrates the order of magnitude of resistance of a 1 cm arterial segment in the systemic circulation (blood viscosity is assumed = 0.04 Poise) :

Table 2.1.

Although calculated hydraulic resistance of the systemic circulation amounts to

approximately 1500 dyne.s.cm^-5 for an adult human in normal physiological conditions, the table illustrates that the major site of resistance to blood flow resides at the level of the arterioles. This statement is confirmed by the fact that the greatest pressure drop in the systemic circulation occurs at the level of the arterioles (figure 2.5). Thus when referring to systemic vascular resistance one merely refers to the resistance offered by the arterioles, because the contribution to hydraulic resistance of the large conductance vessels is small.

Internal radius Internal radius in cm R in dynes.s.cm^-5

Aorta 12 mm 1.2 0.05

medium sized artery 2 mm 0.2 63.7

Arteriole 15 µ 0.0015 (= 1.5 x 10^-3) 2.10⁺¹⁰