Perception, Perfusion & Posture

Billy L Luu

Ph.D. Thesis University of New South Wales 2010

THE UNIVERSITY OF NEW SOUTH WALES Thesis/Dissertation Sheet

Surname or Family name: LUU

First name: BILLY Other name/s: LIANG

Abbreviation for degree as given in the University calendar: PhD

School: MEDICAL SCIENCES Faculty: MEDICINE

Title: PERCEPTION, PERFUSION & POSTURE

Upright posture is a fundamental and critical human behaviour that places unique demands on the brain for motor and cardiovascular control. This thesis examines broad aspects of sensorimotor and cardiovascular function to examine their interdependence, with particular emphasis on the physiological processes operating during standing. How we perceive the forces our muscles exert was investigated by contralateral weight matching in the upper limb. The muscles on one side were weakened to about half their strength by fatigue and by paralysis with curare. In two subjects with dense large-diameter neuropathy, fatigue made lifted objects feel twice as heavy but in normal subjects it made objects feel lighter. Complete paralysis and recovery to half strength also made lifted objects feel lighter. Taken together, these results show that although a signal within the brain is available, this is not how muscle force is normally perceived. Instead, we use a signal that is largely peripheral reafference, which includes a dominant contribution from muscle spindles. Unlike perception of force when lifting weights, it is shown that we perceive only about 15% of the force applied by the legs to balance the body when standing. A series of experiments show that half of the force is provided by passive mechanics and most of the active muscle force is produced by a sub- cortical motor drive that does not give rise to a sense of muscular force through central efference copy or reafference. In addition to feeding perceptual centres of the brain, an efference copy of the motor drive also feeds cardiovascular centres to generate a pressor response during volitional contractions, the effect of sub-cortical motor output on the central pressor response was determined. It is shown that the postural leg muscles depend strongly on their high perfusion pressure during orthostatic posture but because the motor drive from the balance system does not generate a central pressor response they do not receive the benefit of increased perfusion pressure to offset the loss of contractility as they fatigue.

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Acknowledgement

Foremost, I would like to express my sincere gratitude to my supervisor Dr Richard Fitzpatrick for giving me the opportunity to undertake my Ph.D studies in his laboratory. His continuous support, encouragement, and guidance throughout my doctoral research has been invaluable in my development and understanding of the principles of scientific research. The experience of working in his laboratory has ingrained in me the desire to broaden my knowledge in all aspects of human physiology. He is a great friend and mentor.

I sincerely thank Dr Jean-Sébastien Blouin for inviting me to spend time in his laboratory in Vancouver. He generously donated his time and expertise to teach me recording and analysis techniques that offered an insight into the technical aspects of scientific research. I am grateful to Dr Brian Day and Dr Jonathan Cole for providing the experimental data presented in my thesis on an extremely rare patient that has peripheral sensory loss.

I would also like to thank my colleagues at Prince of Wales Medical Research Institute for their participation in my experiments and for creating a wonderful environment for learning. In particular, I would like to acknowledge the other members of my laboratory, Annie Butler and Rebecca St George, for their encouragement, insightful discussions, and for making this an enjoyable experience. To Mr Hilary Carter, I am grateful for the numerous pieces of equipment that were developed for my research.

To my family, I thank them for their support and understanding throughout my studies and for their encouragment to pursue my interests in physiology.

ORIGINALITY STATEMENT ‘I hereby declare that this submission is my own work and to the best of my knowledge it contains no materials previously published or written by another person, or substantial proportions of material which have been accepted for the award of any other degree or diploma at UNSW or any other educational institution, except where due acknowledgement is made in the thesis. Any contribution made to the research by others, with whom I have worked at UNSW or elsewhere, is explicitly acknowledged in the thesis. I also declare that the intellectual content of this thesis is the product of my own work, except to the extent that assistance from others in the project's design and conception or in style, presentation and linguistic expression is acknowledged.’

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COPYRIGHT STATEMENT ‘I hereby grant the University of New South Wales or its agents the right to archive and to make available my thesis or dissertation in whole or part in the University libraries in all forms of media, now or here after known, subject to the provisions of the Copyright Act 1968. I retain all proprietary rights, such as patent rights. I also retain the right to use in future works (such as articles or books) all or part of this thesis or dissertation. I also authorise University Microfilms to use the 350 word abstract of my thesis in Dissertation Abstract International. I have either used no substantial portions of copyright material in my thesis or I have obtained permission to use copyright material; where permission has not been granted I have applied/will apply for a partial restriction of the digital copy of my thesis or dissertation.'

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Abstract

Upright posture is a fundamental and critical human behaviour that places unique demands on the brain for motor and cardiovascular control. This thesis examines broad aspects of sensorimotor and cardiovascular function to examine their interdependence, with particular emphasis on the physiological processes operating during standing.

How we perceive the forces our muscles exert was investigated by contralateral weight matching in the upper limb. The muscles on one side were weakened to about half their strength by fatigue and by paralysis with curare. In two subjects with dense large-diameter neuropathy, fatigue made lifted objects feel twice as heavy but in normal subjects it made objects feel lighter. Complete paralysis and recovery to half strength also made lifted objects feel lighter. Taken together, these results show that although a signal within the brain is available, this is not how muscle force is normally perceived. Instead, we use a signal that is largely peripheral reafference, which includes a dominant contribution from muscle spindles.

Unlike perception of force when lifting weights, it is shown that we perceive only about 15% of the force applied by the legs to balance the body when standing. A series of experiments show that half of the force is provided by passive mechanics and most of the active muscle force is produced by a sub-cortical motor drive that does not give rise to a sense of muscular force through central efference copy or reafference.

In addition to feeding perceptual centres of the brain, an efference copy of the motor drive also feeds cardiovascular centres to generate a pressor response during volitional contractions, the effect of sub-cortical motor output on the central pressor response was determined. It is shown that the postural leg muscles depend strongly on their high perfusion pressure during orthostatic posture but because the motor drive from the balance system does not generate a central pressor response they do not receive the benefit of increased perfusion pressure to offset the loss of contractility as they fatigue.

ii

Contents

Acknowledgement Originality, copyright and authenticity statements Abstract i Contents ii Preface iii Presentations from this thesis iv

Chapter 1 Introduction: The physiology of human posture 1 Chapter 2 On the sense of force and weight 30 Chapter 3 The sensation of standing 51 Chapter 4 Perfusion and lower-limb muscle function 71 Chapter 5 Cardiovascular control during standing 91 Chapter 6 Conclusions and speculations 102

References 112

iii

Preface

This thesis brings together two broad areas of physiology to study how people stand. It focuses on the neural signals that control muscle and cardiovascular activities during human standing. For most motor actions these two physiological systems are tightly coupled, which is not surprising given that the molecules carried by the blood are fundamental to muscle function. The approach throughout this thesis is to investigate the cortical control processes that are common to both of these systems to determine how they are coordinated during standing. Chapter 1 provides an overview of the current ideas in the neural control of standing as well as a review of the literature that focuses on the two aspects of motor control physiology that are examined in this thesis. Each study is presented as a separate chapter in a style that is suitable for publication in a scientific journal. I propose to submit each study for publication in a form similar to the presentations here. The diverse range of psychophysical and physiological experimental approaches used in this thesis meant that some experiments were performed in other laboratories. Chapters 2 – 5 involved psychophysical measurements, muscle stimulation, and physiological recording of blood pressure that were conducted at the Prince of Wales Medical Research Institute, Sydney. In Chapter 2 which investigated perceptions of muscle force, the design required two patients with a very rare sensory disorder who could not be brought to the laboratory. They were tested at the University College London and in rural NSW by Richard Fitzpatrick using the equipment and protocols that I developed in Sydney. An experiment in Chapter 4 required measurements of electroencephalographic activity to assess the contribution of the cortex in the control of standing. I conducted this experiment at the University of British Columbia, Vancouver in collaboration with Dr Jean-Sébastien Blouin. All experimental procedures were approved by the Human Research Ethics Committee of the University of New South Wales and by the research ethics committees from each institution mentioned above. All subjects provided written informed consent before participating. iv

Presentations from this thesis

Luu, BL and Fitzpatrick, RC. 2005. Effects of muscle fatigue on the perceived heaviness of lifted weights. Proceedings of the 25th Annual Meeting of the Australian Neuroscience Society, Perth, Australia, P187.

Luu, BL and Fitzpatrick, RC. 2007. Perceived force of contraction of postural muscles during standing. IBRO World Congress of Neuroscience Satellite Meeting: Motor Control at the Top End, Darwin, Australia, P84.

Luu, BL and Fitzpatrick, RC. 2007. Effects of arterial perfusion pressure on the performance of human leg muscles. Proceedings of the 37th Annual Meeting of the Society for Neuroscience, San Diego, USA, 727.2.

Luu, BL and Fitzpatrick, RC. 2008. The central pressor response during standing. Australian Society for Medical Research (New South Wales) 16th Annual Scientific Meeting, Sydney, Australia, 27.

Luu, BL and Fitzpatrick, RC. 2008. The perceived lightness of standing. Australian Research Council Network in Human Communication Science Workshop: Speech, Perception and Action, Sydney, Australia, 21.

Luu, BL and Fitzpatrick, RC. 2008. The regulation of blood pressure during standing and the importance of hydrostatic load. Proceedings of the 38th Annual Meeting of the Society for Neuroscience, Washington, USA, 676.5.

Luu, BL, Blouin, J-S and Fitzpatrick, RC. 2008. Coherence between cortical and muscular activity during balance and non-balance tasks. 3rd Annual Brain Sciences University of New South Wales Symposium, Sydney, Australia, 30.

Luu, BL, Blouin, J-S and Fitzpatrick, RC. 2009. Changes in cortical activity and perceived muscle force during balance and non-balance tasks. Proceedings of the 39th Annual Meeting of the Society for Neuroscience, Chicago, USA, 662.3.

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Chapter 1

The physiology of human posture

The evolution of our unique upright posture, standing on two legs with the body extended, has required a range of physiological problems to be solved. The ease of standing belies the complexity of the control processes that keep the body balanced. Many different muscles need to be activated precisely or the unstable segments of the body will topple. To coordinate this, the brain relies on information from different sensory receptors to provide information about the gravitational orientation of the body. Our unstable posture means that the muscles that balance the body must act continuously and efficiently. This requires a cardiovascular control process that can operate across a wide range of pressures largely created by the vertical alignment of the body. Blood flow to muscles must be regulated to support the motor activity while protecting blood flow to the brain and other vital organs. The purpose of this thesis is to explore the interactions that coordinate the motor control and cardiovascular control of human standing. Human standing

During human standing, most of the weight of the body is borne by the articulated skeleton and distributed between the two legs. If our upright posture had the body’s centre of gravity aligned with the supporting joints, joint reaction forces and tensile forces in the ligaments could maintain standing almost entirely through passive forces with very little active contribution from the musculature. However, the natural standing posture places the body’s centre of gravity vertically in front of the knee and ankle joints and approximately halfway between the heel and the heads of the metatarsals (Hellebrandt et al., 1938; Fox & Young, 1954; Smith, 1957; Woodhull et al., 1985). Consequently, the body would fall forwards about the ankles if it were not opposed by active contractions of the calf muscles. This arrangement requires continuous activity in the triceps surae muscles to produce an equal but opposite force that pulls the body INTRODUCTION: POSTURE 2 

backwards against gravity. Indeed, electromyographic recordings in the triceps surae indicate a continuous drive to the soleus muscle during normal standing, with more irregular and intermittent bursts of activity observed in the gastrocnemius muscle (Joseph & Nightingale, 1952; Smith, 1954; Portnoy & Morin, 1956). Passive components of the ankle joint, such as ligaments, tendons, and surrounding tissue, contribute a smaller fraction of the plantar force that opposes gravity.

The inverted pendulum The angular motion of the body during standing is considered to approximate the behaviour of an inverted pendulum pivoted at the ankle joints (Smith, 1957; Gurfinkel & Osovets, 1972; Fitzpatrick et al., 1992). As the body sways forward, the ankle torque required to prevent the body from falling can be determined from the angle of forward lean of the body (Smith, 1957). The physics of an inverted pendulum dictates that increasing the angle of forward lean increases the torque required to resist the toppling force of gravity. Conversely, positioning the pendulum’s centre of gravity over the pivot reduces the torque required to stabilise the pendulum. A detailed mathematical description of the inverted pendulum model of standing was provided by Gurfinkel and Osovets (1972) and has since been shown experimentally to be a good approximation of the relationship between ankle torque and body sway during standing (Fitzpatrick et al., 1992; Winter et al., 1998; Loram et al., 2001).

The unstable equilibrium of standing means that any deviation away from the “set point” of balance further increases the destabilising torque. In our upright posture we have a small base of support and relatively high centre of gravity, which is located approximately at the level of the sacral vertebrae (55% of body height) (Croskey et al., 1922; Cotton, 1931; Hellebrandt et al., 1938). Accordingly, the centre of gravity fluctuates around an apparent equilibrium point. Whether we stand relaxed or attempt to stand as still as possible, the body is constantly moving. This natural movement of the body is termed postural sway. With the feet planted firmly on the ground, movement of the body occurs mostly in the anterior-posterior direction as the body sways in an arc around the ankles (Hellebrandt et al., 1938; Orma, 1957; Valk-Fai, 1973; Britton et al., 1993).

INTRODUCTION: POSTURE 3 

The equivalent body The ankle torque generated by muscle contractions against the load of the human body can be replicated with an equivalent mechanical body. By having subjects move an external inverted pendulum back and forth by modulating the activity in the plantar flexors, Fitzpatrick et al. (1992) demonstrated that the slope of the ankle torque versus angle relationship (i.e. load stiffness) of the external body can be matched accurately to the physical load properties of the human body during standing. In their setup, subjects were strapped to a rigid post that supported them in an upright position from behind. A weight approximating the individual subject’s weight and at approximately the height of the subject’s centre of mass was attached to a bar that was coupled to foot plates that rotated about an axis through the ankle joints. Dorsiflexing the ankles to rotate the foot plates caused the vertical bar to lean backwards and apply an increasing torque to the ankles in the same way that ankle torque increases when the person leans forward. Therefore, the mechanical load could be regarded as an 'equivalent person' as it imposes the same torque, moment of inertia, and normal ground reaction force at the ankles and feet as the subject's own body.

The equivalent body technique simulates the mechanical aspects of normal standing but excludes vestibular inputs from directing the appropriate changes in ankle torque to control balance. It was reasoned that immobilising the subject’s head and body, with only their feet used to balance the equivalent body, prevented changes in vestibular afferent information (Fitzpatrick et al., 1992). This was later confirmed using galvanic vestibular stimulation. The galvanic stimulus that produces a stereotyped vestibular evoked muscle reflex and postural response during standing (Coates, 1973; Nashner & Wolfson, 1974; Lund & Broberg, 1983; Britton et al., 1993) was absent when subjects balanced the equivalent mechanical body (Fitzpatrick et al., 1994a). Therefore, the sensory contribution to postural control could be limited to the proprioceptive system by eliminating visual feedback and preventing movement of the body. This equivalent task has subsequently been used in many studies to discern the underlying mechanisms involved in balancing the real body during standing (Fitzpatrick & McCloskey, 1994; Fitzpatrick et al., 1994b; Fitzpatrick et al., 1996a; Winter et al., 1998; Loram et al., 2001). INTRODUCTION: POSTURE 4 

The inverted pendulum model of standing (Gurfinkel & Osovets, 1972) and the equivalent task described by Fitzpatrick et al. (1992) are used with the following caveats: the body is considered as a rigid inverted pendulum rather than a segmented pendulum and movement is restricted to the ankles in the sagittal plane. Using a rigid- body model does not alter the gross dynamics of postural control but does result in an increase in sway size by approximately one-third, indicating an overall stabilising effect of the upper body segments when appropriately controlled (Fitzpatrick et al., 1994b).

Neural control of standing Earlier views of the neuromuscular control of human posture stemmed from the classical work by Sherrington (1906) on decerebrate cat preparations. Sherrington systematically described the effects of a transection at different levels of the neuraxis on muscle tone. A complete lesion of the spinal cord resulted in a loss of tonic background muscle activity and the animal’s ability to stand. A transection of the brainstem that separated the cerebrum, but preserved pontine, medullary and spinal pathways, resulted in an exaggerated increase of muscle tone in the extensor muscles. This produced a rigid standing posture that was referred to as decerebrate rigidity. Sherrington interpreted the manifestations of decerebrate rigidity as a release of inhibitory cortical influences on the excitatory actions of the pontine reticular formation and Dieter’s nucleus (the lateral vestibular nucleus) on tonic muscle activity. Consequently, the vestibular sense organs, which project to these brainstem centres, were associated with the tonic signal to the extensor muscles during standing.

A phenomenon similar to decerebrate rigidity is also observed clinically. Lesions to the brainstem produce a rigid and extended body posture with pronounced extensor muscle tone in the neck, trunk, and upper and lower limbs (Wilson, 1920; Walshe, 1923; Davis, 1925). These observations suggest that brain stem centres exert a tonic influence on extensor muscle tone, and it is now thought that this is mediated through the bistable property of spinal motoneurones (Kiehn & Eken, 1997; Gorassini et al., 1998). However, human brain stem lesions are caused predominantly by trauma and ischaemia and do not create the complete and precise lesions attained in animal experiments. Human ‘decerebrate rigidity’ is usually accompanied by unconsciousness, which limits observations to the supine posture and not during the process of controlling upright posture. Furthermore, increases in extensor muscle tone are indiscriminate as INTRODUCTION: POSTURE 5 

extensor rigidity extends to the muscles of the upper limb that are not typically involved in postural control. Regardless of these differences, it has been suggested that a tonic drive to the postural muscles exists in human standing (Hellebrandt & Franseen, 1943; Martin, 1967; Winter et al., 1998; Gurfinkel et al., 2006). However, difficulties in isolating and recording activity from the brain stem in conscious human subjects have resulted in reservations to whether tonic activity is mediated by central processes or driven by sensory afferents.

Tonic muscle activity is recognised as only one aspect of postural control: the maintenance of upright posture against static disturbances. In particular, static postural control is thought to establish the background muscle activity that determines the basic standing posture against the constant force of gravity. The stabilisation of posture in response to random disturbances involves a separate dynamic control process that is characterised by reactive reflex adjustments in body position. Essentially, sensory organs detect changes in body position and evoke an appropriate muscle response to restore the body to its original standing position.

Several sensory systems can contribute to the dynamic control of upright posture. As such, numerous studies have been conducted to identify which sensory cues are relied upon during standing. This typically involves excluding or enhancing a particular sensory cue to determine the reliance placed on that input. These approaches have revealed that visual signals of self movement (Edwards, 1946; Paulus et al., 1984; Paulus et al., 1989; Fitzpatrick et al., 1994b), vestibular signals of gravitational orientation (Bussel et al., 1980; Allum & Pfaltz, 1985; Keshner et al., 1987), proprioceptive signals of ankle movement (Aggashyan et al., 1973; Mauritz & Dietz, 1980; Diener et al., 1984; Magnusson et al., 1990; Fitzpatrick et al., 1994b) and contact-force signals from mechanoreceptors in the feet (Magnusson et al., 1990; Fitzpatrick et al., 1994b) all play important roles in balance control. Standing is most stable when all sensory inputs are available.

The attraction of a dynamic postural control system that is mediated by sensory reflex reactions to postural disturbances has been considered. Hellebrandt (1938) advocated the stretch reflex strategy as a means of postural stabilisation. It was reasoned that displacements in the body’s centre of gravity stretched or shortened the calf INTRODUCTION: POSTURE 6 

muscles as the body swayed, producing a muscle spindle input to stretch reflexes. Destabilising events that increased postural sway would stretch the calf muscles and evoke a compensatory reflex contraction to restore the body to its original position. These reactive postural adjustments have been shown to be useful in large and unexpected destabilising perturbations (Nashner, 1976; Dietz et al., 1989; Nardone et al., 1990). However, the effectiveness of this strategy during normal quiet standing was questioned by studies that indicated that the range of motion at the ankles during natural sway was not large enough to elicit a stretch reflex response (Kelton & Wright, 1949; Gurfinkel et al., 1974). There is also evidence that the H-reflex, the electrical analogue of the short-latency stretch reflex, is inhibited presynaptically during standing (Katz et al., 1988; Koceja et al., 1993), suggesting a decreased reliance on these simpler neural processes. Other studies have expressed concerns about feedback instability caused by the large delays in neural transmission (Fitzpatrick et al., 1996a; Morasso & Schieppati, 1999).

Based on their estimates of muscle stiffness and derivations of ankle torque from the inverted pendulum model of standing, Gurfinkel et al. (1974) and Nashner (1976) suggested that the inherent stiffness properties of muscles could contribute to the initial compensatory response to small deviations in posture. This view was developed further by Winter et al. (1998; 2001) to include the assumption that muscles behave like springs in response to balance perturbations. Winter et al. proposed that the central nervous system sets the ankle stiffness at an appropriately high level to stand by maintaining a constant muscle tone, while the muscle itself functions like a spring to provide an instantaneous corrective response to imposed changes in posture. This spring-stiffness model of standing was based on the premise that the ankle joint can be made sufficiently stiff to support the load of the body. While this assertion may be true about the stiffness properties of active muscles, it does not translate to the total stiffness of the ankle joint which is provided by the combination of muscle short-range and thixotropic effects and the elastic stiffness of tendons and surrounding tissues. Recent findings (Loram & Lakie, 2002; Morasso & Sanguineti, 2002; Casadio et al., 2005) indicate that the intrinsic ankle stiffness lies between 65 – 91% of the total ankle stiffness required to support the body’s load during standing. It was argued that the series linkage between the calf muscles and the compliant Achilles’ tendon limits the stiffness of the ankle joint to below that INTRODUCTION: POSTURE 7 

necessary to stand (Loram & Lakie, 2002). Essentially, the ankle joint is unable to increase ankle stiffness sufficiently, despite increases in calf muscle activity, as the compliant Achilles’ tendon inevitably stretches under the body’s load to a point where the body eventually topples. An important implication of this model is that there is no level of tonic muscle activity that will maintain upright posture and therefore some level of modulated drive is critical.

More recently, Loram et al. (2004, 2005a, b) demonstrated an unexpected behaviour in the dependent leg muscles during standing. Using ultrasound imaging, they noticed that during forward sway the gross length of the calf muscle and tendon unit increased while the length of the muscle fascicles shortened. This is possible because of the compliant Achilles’ tendon (Loram & Lakie, 2002). The authors refer to this unorthodox muscle behaviour as paradoxical muscle movements, a characteristic that appears to be intrinsic to balancing the inverted load of the upright human body. These paradoxical muscle movements occur periodically, on average 2.8 times per unidirectional sway of the body (Loram et al., 2005b), and may reflect the means by which the central nervous system resolves the issue of balancing a load that is stiffer than the Achilles’ tendon. This cyclic pattern of muscle activity in which bursts of electromyographic activity in the calf muscles precede paradoxical shortening produces repeated ballistic, catch-and-throw movements of the body that cannot be explained by simple stretch reflexes alone (Loram et al., 2005b). Rather, this paradoxical muscle behaviour suggests that the increase in ankle torque recorded during forward sway is generated mostly by active muscle contractions and not from mechanical stretching or spring-like properties of the muscles. The neural mechanisms that generate these paradoxical movements have not been established, although feed-forward or anticipatory control by the central nervous system is proposed as the basis of standing balance (Loram et al., 2005b). INTRODUCTION: PERCEPTION 8

Perceptions of force

Knowledge of the distribution of muscle forces in the body is essential for planning and executing precise motor actions. The sense of force exerted provides information about the adequacy of the muscle contraction for any task. For this purpose, our perception of force is explicably linked to the control of the musculoskeletal system. The means by which we perceive force is complex and a rich history of debate over the last century has seen changing views on the neural processes that underlie this sensation. The information upon which this perception is based could come from two potentially conflicting sources; a signal of tension from the periphery or a centrally-generated sense of effort that is related to the motor drive to the muscles. A scheme of the neural inputs that contribute to conscious perception, and most likely to the perception of force, is illustrated in Figure 1.1.

Central contributions to force perception The sense of effort was proposed by philosopher George Henry Lewes (1879), who suggested that efferent motor commands within the central nervous system might somehow contribute to the sense of effort involved in movement. His theory was based on observations that passive movement is largely devoid of sensation compared with volitional actions. Lewes’ proposal now forms the contemporary view that the sense of effort is derived from a corollary discharge that is related to the centrally-generated voluntary motor command (McCloskey et al., 1974). That is, the corollary discharge is a signal existing at some level within the central nervous system as a copy of the efferent signal for muscle contraction (Sperry, 1950). Sperry hypothesised that the corollary discharge acted via perceptual centres and interacted with afferent signals to guide motor behaviour.

The term efference copy, which was central to a different concept of perception (von Holst & Mittelstaedt, 1950, 1973), is often used interchangeably with corollary discharge, and despite them historically having distinct functional purposes (Donaldson, 2000) both terms imply that the central nervous system preserves a copy of the central motor command. Von Holst and Mittelstaedt’s reafference principle proposes that a central efference copy of the motor action is used to compare with the total afferent signal to separate a reafference signal component that represents the consequences of self- INTRODUCTION: PERCEPTION 9

generated motor actions, and a exafference signal component that is the consequence of external influences. In this sense, the efference copy provides an expectation of the afferent signals generated by our own actions and defines the circumstances in which the afferent signals arise.

Figure 1.1 Central and peripheral mechanisms involved in force perception Afferent fibres from skin (MC, PC, MD, RE), joint (JR) and muscle (MS, GTO) known to contribute to conscious perception are displayed in orange. Efferent outflow, displayed in blue, is a derivative of the cortical command. Corollary discharge that could contribute to conscious perception is believed to originate directly from the cortical command.

With normal healthy muscles, subjects can match the force applied during a muscle contraction and the perceived force can be matched with reasonable accuracy (approximately 20% coefficient of variation) by weight matching with the contralateral limb (McCloskey et al., 1974; Gandevia & Kilbreath, 1990). Von Helmholtz (1867) proposed that sensations of innervation during voluntary muscle contractions, a term INTRODUCTION: PERCEPTION 10

that had similar connotations to the sense of effort, provided sensations attributable to the muscles. This has since been borne out by many studies reporting that subjects prefer to use their sense of effort in judging the perceived force or heaviness of a lifted weight (McCloskey et al., 1974; Gandevia & McCloskey, 1977a; Jones & Hunter, 1983a, b; Cafarelli, 1988; Carson et al., 2002; Proske et al., 2004). To demonstrate this central-command effect, these studies have relied on subjects’ perceptions of force under conditions of relative muscle weakness.

Whenever a muscle is weakened, by paresis, fatigue or neurological disorder, the perception of force becomes distorted. Patients with paresis through upper motor neurone loss feel weights to be heavier on the affected side (Gandevia & McCloskey, 1977a). When paresis is induced experimentally, by infusing a neuromuscular blocking agent, the force generated by the affected muscle is perceived to be greater than under normal conditions (Gandevia & McCloskey, 1977a). Similarly, patients suffering myasthenia, a condition that blocks neuromuscular transmission, frequently complain of the heaviness of their limbs. An explanation for these behaviours is that to produce the same force output, a weakened muscle requires a greater motoneuron drive than a healthy muscle. Consequently, the increase in motoneuron drive and effort translates into an increase in perceived force.

Similar effects are observed with muscle fatigue. Numerous studies have shown that muscle weakness induced by prolonged fatiguing contractions alters the perception of force (McCloskey et al., 1974; Gandevia & McCloskey, 1978; Jones & Hunter, 1983a, b; Cafarelli & Layton-Wood, 1986; Burgess & Jones, 1997; Proske et al., 2004). As the efficiency of muscle contraction decreases as fatigue develops over time, an increase in motoneuron drive is required to maintain a constant force. Thus, the corresponding increase in effort results in the same contraction force feeling stronger (Gandevia & McCloskey, 1977a).

Further evidence that neural drive to muscles is associated with the perception of force is obtained from a study on a single subject that has a complete large sensory-fibre peripheral neuropathy below neck level. The subject had lost the sense of limb movement, limb position, and cutaneous touch but had preserved motor functions (Cole & Sedgwick, 1992; Cole et al., 1995). Despite the loss of proprioceptive abilities, the INTRODUCTION: PERCEPTION 11

subject was able to match the force required to lift a weight with the contralateral hand, albeit with a directional bias that resulted in a consistent overestimation of contraction force (Cole & Sedgwick, 1992).

Peripheral contributions to force perception Unlike the sense of effort, the sense of tension has a peripheral connotation. It refers to a sense of tension or force within muscles and tendons. Sensations of muscular tension arise from the sensory receptors of the peripheral nervous system, which are classified as skin, joint, or intramuscular receptors (Figure 1.1). Each type of receptor conveys information that contributes to different aspects of perception. The receptors of all classes are innervated by neurones in the dorsal column nuclei with fast-conducting group I or group II afferents. Afferent pathways ascend in the dorsal columns of the spinal cord and are relayed through the cuneate or gracile nuclei of the medulla and the posterior thalamic nuclei before reaching the somatosensory cortex.

If the perception of force was based on peripheral mechanisms, perceived force would be unchanged for a weakened muscle as the actual muscle tension is unchanged. As many previous studies have shown, perceived force increases with muscle weakness (McCloskey et al., 1974; Gandevia & McCloskey, 1976b, 1977a, b; Jones & Hunter, 1983a, b) indicating that under normal conditions when a muscle is not weakened, peripheral receptors appear not to be used in the judgement of muscle force. However, there is evidence suggesting that signals from the periphery also contribute to the perception of force. First, in the absence of any centrally generated input to the muscle during reflex contractions from tendon vibration (Brodie & Ross, 1984) and electrical stimulation (Waller, 1891), discrimination of weights is still possible, albeit with poorer performance. Second, anaesthesia of the hand makes weights lifted by the biceps muscle feel heavier (Gandevia & McCloskey, 1976b) and similarly anaesthesia of the thumb increases perceived heaviness when lifting a weight by thumb flexion (Gandevia & McCloskey, 1977a; Marsden et al., 1978, 1979; Gandevia et al., 1980). These studies indicate that peripheral input of tension from joint and skin afferents influence perceptions of muscle force.

Joint receptors. Sensory receptors within joint capsules and surrounding ligaments have traditionally been associated with the detection of movement (Gandevia & McCloskey, INTRODUCTION: PERCEPTION 12

1976a; Ferrell et al., 1987) as they respond to stretch in a direction specific manner (Macefield et al., 1990). These receptors respond with large changes in firing rates for movements at the angular extremes of joint rotation towards hyperflexion and hyperextension and show only small changes in firing rate for most of the range of joint movement (Goodwin et al., 1997). In addition, blocking joint afferents results in systematic errors at the extremes of joint movement (Ferrell et al., 1987). Joint receptors, therefore, are most suitable for signalling the limits of joint rotation and so far have not been implicated in the appreciation of muscle force.

Skin receptors. There are many types of sensory receptors in the skin, but those of concern here are the group of mechanoreceptors that respond to touch and pressure. Skin mechanoreceptors are classified according to their rate of adaptation to sustained pressure (fast or slowly adapting) and their receptive field sizes (Type I – small, well- defined field; Type II – large, poorly-defined field) (Vallbo & Johansson, 1984). In particular, slowly-adapting type II afferents of the skin, also known as Ruffini endings, appear to be suited to provide information about pressure against the skin as they respond vigorously to indentation and they can signal contact force (Goodwin et al., 1997). The rapidly-adapting type skin receptors, Meissner’s and Pacinian corpuscles, respond transiently without grading to force. Thus, they appear specialised to provide information about the timing and location of skin contact with single or repeated stimuli such as vibration (Vallbo et al., 1979; Macefield et al., 1990).

Accuracy in discriminating contact force and pressure by skin mechanoreceptors varies for different parts of the body. The fingertips have a higher degree of accuracy compared with the rest of the body since they have very low thresholds in response to applied touch and pressure, are very densely distributed and have much smaller receptive fields (Vallbo et al., 1979; Vallbo et al., 1984). Nevertheless, the same types of skin receptors are found in the upper (Johansson & Vallbo, 1983) and lower limbs (Vallbo & Hagbarth, 1968; Ribot-Ciscar et al., 1989) as well as in the soles of the feet (Kennedy & Inglis, 2002).

Intramuscular receptors. It is perhaps surprising that cutaneous receptors have been shown to affect the sensation of tension when the most apparent source of peripheral information could be obtained from sensory receptors within the muscle itself. There INTRODUCTION: PERCEPTION 13

are two types of peripheral receptors present in muscles: muscle spindles, which are aligned parallel to the extrafusal force-producing muscle fibres, and tendon organs, which are orientated in series. The number and size of each type varies considerably with each muscle.

Muscle spindles do not appear to contribute to the perception of force. Despite progressive increases in muscle electrical activity during a maintained constant-force sub-maximal voluntary contraction, the discharge rate of muscle spindles declines (Macefield et al., 1991). At moderate to high contractile forces, the cumulative rate of increase in firing diminishes as muscle spindles appear to approach saturation (Vallbo, 1974b; Wilson et al., 1997). These two properties of muscle spindles make them an impractical source of information about muscle force. Furthermore, when muscle spindles are excited by tendon vibration, perception of muscle force is diminished (Jones & Hunter, 1983b). This is opposite to that expected if they provided a signal of muscle tension and in line with some of the contraction being produced through the tonic vibration reflex operating at sub-perceptual levels (Marsden et al., 1969), further supporting the central force-perception hypothesis.

In contrast, the musculotendinous sensory corpuscle, described by Camillo Golgi in 1878 and which now bears his name, reliably signals the local force generated within a muscle by contraction or stretch. Golgi tendon organs are particularly sensitive to longitudinal stretch and are able to monitor tension developed in a specific muscle. Individually, they are more sensitive to tension developed by muscle contraction than by passive stretch, probably because of a focussing of the contractile force to the region of the tendon organ (Gregory & Proske, 1979). They respond to contractions of more than one motor unit and are therefore specialised to signal the overall load or tension in a muscle, independent of length (Gregory & Proske, 1975, 1979). The information obtained from Golgi tendon organs is thought to provide an ideal signal to use for judgments of muscle force (Crago et al., 1982; Jami, 1992).

Of the potential sources of information available to base our estimates of muscle force on, the consensus is that we prefer to perceive force by attending to the sense of effort that is associated with the central motor commands to the muscles. This view developed from the demonstration by McCloskey et al. (1974) that when the INTRODUCTION: PERCEPTION 14

excitability of the motoneuron pool was artificially altered with muscle tendon vibration, the subject’s perception of force was guided by their effort of contraction rather than the actual tension developed within the muscle. However, there is a problem with the sense of effort as the exclusive determinant of perceived force. Simply put, if a muscle is rendered half as strong, a lifted object does not feel twice as heavy. In the flexors of the index finger, perceived force increases by only about 75% when the muscles are weakened to half their strength (Gandevia & McCloskey, 1977a, b). There are two possibilities for this discrepancy: either the central command does not scale linearly or the contributions from central and peripheral mechanisms to the perception of force are not mutually exclusive as over-estimation of weights, due to increases in effort, would be much greater than those reported (McCloskey et al., 1974; Gandevia & McCloskey, 1977a; Jones & Hunter, 1983b). On the other hand, if tension was only monitored by peripheral mechanisms, the perception of force would essentially remain the same. This suggests that the actual perception of force during muscle fatigue lies in between these two extremes.

Force perception during standing Personal experience tells us that we do not require a substantial amount of effort to stand. This is somewhat surprising considering the plantar flexors are continuously active when balancing the body. Whether this perceived ease of standing is a true reflection of the muscular force generated by the plantar flexors is not known since studies conducted on force perception have been carried out exclusively in non-load bearing tasks with the subject seated (McCloskey et al., 1974; Gandevia & McCloskey, 1976b, 1977a, b; Jones & Hunter, 1983a, b; Cafarelli & Layton-Wood, 1986; Cafarelli, 1988; Gandevia & Kilbreath, 1990; Tremblay et al., 2001). It is possible that the perceived ease of standing is simply because the force that the feet exert on the floor to prevent the body from falling forwards – the applied plantar force – requires only a weak contraction. It is also possible that in the upright posture, the normal ground reaction force due to the weight of the body, which is separate from the applied plantar force but also acts through the feet, influences, or perhaps masks, the perceived force exerted by the plantar flexors during standing.

INTRODUCTION: PERFUSION 15

Cardiovascular aspects of human standing

The regulation of the cardiovascular system, comprising the heart and circulation, logically must be an integral part of human motor control. Muscle performance depends on a continuous and controlled movement of blood to where it is needed as the circulating blood transports oxygen and nutrients to active tissues and removes the build up of by- products and carbon dioxide from tissue metabolism. The circulatory system, first described by William Harvey in 1628, essentially represents a closed system where blood from the heart travels through the arterial circulation to reach the network of capillaries, where the exchange of gases, molecules and fluids takes place, and returns to the heart via the venous circulation. As a result, the volume of blood that the heart can pump is limited by the volume of blood that returns from the venous circulation (Guyton et al., 1957; Guyton et al., 1962). The rate at which blood circulates is determined by the pressure and the resistance of the vessels since, like all fluids, it flows from the regions of higher pressure to regions of lower pressure.

Blood pressure Blood pressure refers to the pressure exerted by the blood against the walls of the blood vessels. It is created by the contractions of the heart, which forces blood into the arterial circulation, and by the resistance of the blood vessels to blood flow. For this reason, blood pressure in the arterial circulation is pulsatile, fluctuating with each rhythmic contraction of the heart. Peak arterial pressure during each cardiac cycle, the systolic pressure, is recorded as the heart pumps blood into the circulation. The minimum arterial pressure, the diastolic pressure, occurs between heart contractions and is a product of the resistance of the arterial vessels and the volume they contain.

As blood moves further away from the heart and large vessels, the driving force of the heart pump dissipates and pulsatile activity and blood pressure progressively decline. In the supine position, blood pressure decreases from a mean value of approximately 95 mmHg when blood leaves the heart to between 17 – 35 mmHg in the capillaries (Hargens et al., 1981) where blood pressure is uniform and blood flow is continuous. In the small veins of the leg, blood pressure is approximately 12 mmHg (Pollack & Wood, 1949). This decreasing pressure gradient from arteries to capillaries to veins facilitates the movement of blood through the circulation. Although the arterial circulation is INTRODUCTION: PERFUSION 16

characterised by a pressure waveform, the force that drives blood through the systemic circulation is more accurately described by the mean arterial blood pressure. Mean pressure can be approximated arithmetically by adding onethird of the pulse pressure (systolic minus diastolic pressure) to the diastolic pressure. Mean arterial pressure is a function of cardiac output and total peripheral vascular resistance. Cardiac output, the product of heart rate and stroke volume, is a computed measure that represents the volume of blood that is expelled by the heart per minute. Venous return is the significant determinant of cardiac output as the normal heart acts as an automatic pump that expels all that is delivered to it (Guyton et al., 1957; Guyton et al., 1962). Changes in either cardiac output or total peripheral resistance alter mean arterial blood pressure. However, increased mean arterial pressure mediated by an increase in cardiac output alone will increase blood flow whereas an increase in mean arterial pressure created by an increase in total peripheral resistance alone will impede blood flow. Physiologically, adjustments in blood pressure are usually achieved by changes in both cardiac output and vascular resistance (Schmidt et al., 1972).

Regulation of blood pressure There are several autoregulatory mechanisms that serve to maintain blood pressure around a long-term operating set-point. These mechanisms can respond rapidly, within one heart beat, to counteract a sudden change in blood pressure or they can act over a period of minutes to hours to oppose prolonged disturbances. The response to abrupt changes in blood pressure is mediated through a neural mechanism that involves the two components of the autonomic nervous system: the sympathetic nervous system which excites the cardiovascular system and the opposing parasympathetic nervous system. For prolonged disturbances in blood pressure, a humoral control mechanism is used that primarily involves the release of vasoactive hormones that act on the renal system. Generally, these mechanisms keep resting mean arterial pressure within the narrow physiological range of 75 – 125 mmHg (Master et al., 1951).

Baroreceptors Slowly-adapting, mechanosensitive baroreceptors are the major compensatory reflex mechanism that responds to short-term fluctuations in blood pressure. They are concentrated within the walls of the aortic arch and the carotid sinuses and respond to circumferential (Landgren, 1952; Angell James, 1971) and longitudinal stretch (Sollmann INTRODUCTION: PERFUSION 17

& Brown, 1912; Angell James, 1971) of the vessel walls over a wide range of intravascular pressures, from approximately 50 to 150 mmHg (Kirchheim, 1976).

The arterial baroreceptor reflex provides tonic excitatory inputs to the central nervous system through the vagus and glossopharyngeal nerves, which terminate in the nucleus tractus solitarius (NTS) of the medulla (Cottle, 1964; Calaresu & Pearce, 1965; Crill & Reis, 1968; Gabriel & Seller, 1970; Culberson & Kimmel, 1972; Lipski et al., 1975; Jordan & Spyer, 1986). Neurons from the NTS convey excitatory baroreceptor signals to the rostral ventrolateral medulla (RVLM) (Ross et al., 1985; Dampney et al., 1987), which directly innervates sympathetic preganglionic neurons in the spinal cord (Ross et al., 1984a; Dampney et al., 1987), and the caudal ventrolateral medulla (CVLM), particularly within the region of the nucleus ambiguus where vagal parasympathetic preganglionic neurons are located (Loewy & Burton, 1978). Additionally, the CVLM has a direct inhibitory projection to the RVLM (Willette et al., 1984; Blessing, 1988). This additional pathway allows the NTS to inhibit sympatho-excitatory activity in the RVLM indirectly as direct inhibitory pathways between the NTS, or the CVLM, and sympathetic preganglionic neurons do not appear to exist (Strack et al., 1989a; Strack et al., 1989b). The nucleus tractus solitarius, along with the discrete medullary regions described above, function as the primary sites of baroreceptor reflex integration (Spyer, 1981, 1982) and forms what has long been considered as the control centre of the cardiovascular system (Alexander, 1946).

An essential element of the baroreceptor reflex is that the RVLM also plays a crucial role in maintaining resting blood pressure. Earlier studies have shown that bilateral lesions or inhibition of RVLM neurons in anaesthetised animals produce profound decreases in resting blood pressure and sympathetic vasomotor activity (Guertzenstein & Silver, 1974; Feldberg & Guertzenstein, 1976). Likewise, an excitation of RVLM neurons promotes wide spread increases in blood pressure and sympathetic outflow (Ross et al., 1984b; McAllen, 1986). These observations have led to the view that RVLM neurons provide a tonic excitatory drive to sympathetic vasomotor neurons that are largely responsible for establishing and maintaining resting arterial blood pressure (Feldberg, 1976; Dampney, 1994; Dampney et al., 2003). INTRODUCTION: PERFUSION 18

The response of the baroreceptor reflex to an elevation in resting arterial blood pressure is illustrated in Figure 1.2. A rise in blood pressure increases tonic baroreceptor activity with the end result being a reduction in cardiac output and total peripheral resistance. These reductions are achieved by two separate means. First, tonic sympatho- excitatory neurons in the RVLM are indirectly suppressed by baroreceptor inputs to the NTS. The net effects of reduced sympatho-excitatory drive are decreased sympathetic vasomotor activity, directed primarily to the resistance vessels of the arterial circulation since they are the major determinant of peripheral vascular resistance, and decreased heart rate and heart contractility. The latter effects on the heart contribute to a reduction in cardiac output. The reflex response to an increase in blood pressure, however, is mediated primarily by an increased parasympathetic (or vagal) outflow to the heart (Schaefer, 1960; Kirchheim, 1976; Spyer, 1981; Thames et al., 1981). Vagal innervations to the heart can act rapidly, within one heart beat (Pickering & Davies, 1973), to lower heart rate considerably and thus reduce cardiac output and arterial blood pressure.

Raising blood pressure is not simply the physiological inverse of lowering blood pressure. A fall in arterial blood pressure is corrected mainly by a sympathetically- mediated increase in total peripheral vascular resistance rather than through changes in cardiac output. Decreased baroreceptor input to the NTS, in response to a fall in blood pressure, allows for an increase in sympathetic vasomotor activity through the disinhibition of tonic sympatho-excitatory neurons in the RVLM (Figure 1.2). Increased vasomotor tone is directed to the resistance vessels of the arterial circulation and is initiated within 5 to 15 seconds (Rowell, 1993). The capacity for an increase in heart rate mediated by sympathetic drive or vagal withdrawal to restore a fall in arterial blood pressure is somewhat limited because a rise in cardiac output is restricted by the opposing effect low arterial blood pressure has on venous return (Guyton et al., 1957).

Cardiovascular adjustments to upright posture The transition from a recumbent to an upright posture creates a problem for the human cardiovascular system. In the recumbent position, most organs and blood vessels are more or less at heart level so that the driving force of arterial blood pressure is relatively uniform from head to toe. Upon standing upright, the head is now positioned above the heart so that to protect the critical cerebral tissues from ischaemia the cardiovascular system must now regulate blood pressure to overcome this effect of gravity that would INTRODUCTION: PERFUSION 19

otherwise reduce cerebral pressure. This includes responding to a redistribution of almost 25 percent of blood volume from the heart and lungs to dependent regions below the heart (Sjostrand, 1952). The profound effect of gravity on the circulation was demonstrated by immersing the upright body in water, bringing arterial pressure and heart rate close to recumbent levels (Hellebrandt & Brogdon, 1938). It was reasoned that the densities of blood and water are similar so placing the body in water effectively minimises the hydrostatic effects of gravity through a redistribution of blood from the extremities.

Figure 1.2 Central pathways subserving the baroreceptor reflex In response to an increase in arterial blood pressure, the cardiovascular control centre (enclosed within the dotted box) induces a reflex increase in parasympathetic activity and a decrease in sympathetic activity to restore arterial blood pressure. Baroreceptor inputs to the nucleus tractus solitarius directly excite the caudal ventrolateral medulla, simultaneously increasing parasympathetic activity to the heart and inhibit tonic sympatho-excitatory neurons in the rostral ventrolateral medulla. The latter effect results in a sympathetically-mediated decrease in cardiac output and total peripheral resistance. Inhibitory connections in the baroreceptor reflex pathway, displayed in red, play a major role in the compensatory response to a rise in arterial blood pressure.

INTRODUCTION: PERFUSION 20

Reflex cardiovascular adjustments for standing There is a distinct difference between the initial blood pressure responses to standing up actively and being tilted passively to an upright posture. A brief, but pronounced drop in mean arterial blood pressure of up to 40 mmHg has been reported during the first 30 seconds of active standing whereas this is normally absent or fairly modest during passive upright tilt (Borst et al., 1984; Sprangers et al., 1991; Tanaka et al., 1996; Heldt et al., 2003). This initial drop in arterial blood pressure during active stance appears to be related to the activation of skeletal muscles rather than an immediate hydrostatic effect on the circulation. Nevertheless, the cardiovascular responses after 1 to 2 minutes of active or passive upright posture are similar and characterised by an increase in mean arterial blood pressure and heart rate and by a decrease in cardiac output (Imholz et al., 1990; Tanaka et al., 1996; Heldt et al., 2003).

Venous pooling explains the decrease in cardiac output on standing. The apparent fall in arterial blood pressure at cervical levels is created in part by the decreased cardiac output and in part by the increased difference in height of the carotid and aortic arch arterial baroreceptors relative to the heart. The lasting changes in mean arterial blood pressure and heart rate reflect the baroreceptor reflex response that offsets this fall in cervical blood pressure. Adopting an upright posture presumably lowers the hydrostatic pressure in the carotid sinus by 15 – 20 mmHg, relative to the aorta. The subsequent decrease in carotid arterial baroreceptor activity induces sympathetically-mediated vasoconstriction in skin, skeletal muscle, renal, and splanchnic regions to increase total vascular resistance and also to prevent blood from accumulating in the periphery. Vasoconstriction is considered to be the main compensatory reflex mechanism preserving mean arterial blood pressure (Tyden, 1977) and it appears to be adequate at least for momentary upright posture. As discussed above, increasing heart rate on standing is ineffective in raising arterial blood pressure as it is unable to increase cardiac output when facing a falling venous return.

Decreased venous return also engages another group of mechanoreceptors located within the walls of the heart and pulmonary circulation. These receptors are referred to as low-pressure or cardiopulmonary baroreceptors and have similar properties to arterial baroreceptors. They respond to changes in transmural pressure by altering peripheral vascular resistance with only minor effects on heart rate and cardiac output (Abboud et INTRODUCTION: PERFUSION 21

al., 1979; Mohanty et al., 1988; Hirsch et al., 1989). In the case of standing, decreased venous return and the corresponding decrease in cardiopulmonary baroreceptor activity can cause a reflex vasoconstriction in the arteries of the skin and skeletal muscle (Johnson et al., 1974).

Despite the efforts of arterial and cardiopulmonary baroreceptor-mediated vasoconstriction, venous return and cardiac output remain compromised during human standing. This is because in man, apart from the skin and splanchnic vessels, vasoconstriction develops entirely in the arterial vessels whereas blood accumulates in the compliant veins of the legs. Clearly, reflex cardiovascular adjustments alone are not sufficient to counteract the blood pressure changes that result from shifts in blood volume during orthostatic posture. If left unabated this can, and often does, ultimately lead to unconsciousness through cerebral ischaemia.

The mechanical pump The importance of movement and skeletal muscle contractions in reducing swelling in the feet had been recognised in as early as the 18th century (referred to in Rowell, 1993). It was hypothesised that rhythmic contractions in the large, lower-limb muscles behaved like a secondary pump in the venous circulation with each contraction actively forcing blood out of the veins and back towards the heart. Indeed, contraction of the muscles can generate intramuscular pressures of up to 300 mmHg (Hill, 1948; Kjellmer, 1964), which is more than sufficient to compress large leg veins completely and overcome the hydrostatic head of pressure that develops during standing. Even with just a gentle contraction by the leg muscles, venous return and cardiac output are restored to recumbent values (Rowell, 1993).

The effectiveness of the muscle pump in promoting venous return depends entirely on working venous valves, as commonly seen in patients with valvular defects (Pollack et al., 1949; Bevegard, 1962). They are a structural adaptation unique to and present throughout the venous vasculature, and they ensure that blood only flows in one direction. Between contractions, venous valves prevent the back flow of blood due to gravity and, with the development of high intramuscular pressures, they prevent blood from being forced back into the arterial circulation. INTRODUCTION: PERFUSION 22

In addition to augmenting venous return, the skeletal muscle pump improves muscle perfusion pressure by creating a greater pressure differential between the arterial and venous circulations. Laughlin (1987) proposed that immediately after a muscle contraction, the muscle pump reduces venous pressure to near zero or even slightly negative pressures. This means that for standing, where arterial blood pressure is approximately 200 mmHg in the lower limb and venous pressure is 120 mmHg (Stegall, 1966), the muscle pump would effectively increase muscle perfusion pressure by 120 mmHg, a measure that is comparable to the driving force of the heart. Since capillaries appear not to be compressed during muscle contractions (Gray et al., 1967), the increase in muscle perfusion pressure would reflect a considerable improvement in muscle blood flow in the legs. It is conceivable then that postural sway, often regarded as an undesirable consequence of imperfect motor control, is an important component in cardiovascular control during upright human standing.

Cardiovascular response to exercise The cardiovascular response to exercise is characterised by measured increases in heart rate, cardiac output, and mean arterial blood pressure. The extent to which these variables change depends on the type of muscle activity (Stebbins & Longhurst, 1995). In general, dynamic contractions evoke large increases in heart rate and cardiac output with only a small effect on mean arterial blood pressure. On the other hand, contractions with a steady force output produce large increases in mean arterial blood pressure with only small increases in heart rate and cardiac output. These differences are presumably due to the decrease in peripheral vascular resistance, and thus diastolic blood pressure, present during dynamic but not static muscle contractions (Kaufman & Hayes, 2002). The neural mechanisms underlying these increases involve both central and peripheral signals acting on the cardiovascular control centres in the brainstem to influence autonomic nervous activity. The central and peripheral mechanisms implicated in the control of the exercise pressor response are illustrated in Figure 1.3.

Central mechanisms The notion of a centrally-generated human pressor response, introduced by Krogh and Lindhard (1913), described a cortical irradiation of the cardiovascular centres by central motor commands. This has evolved into the current view that, not unlike our sense of effort, corollary discharges arising from the central motor command contribute to the INTRODUCTION: PERFUSION 23

initial rise in heart rate and blood pressure during exercise (Figure 1.3). Early investigations by Lind and McNicol (1967a) provided support for a centrally-generated pressor response. They demonstrated that increases in heart rate and blood pressure corresponded to the relative intensity of the motor command rather than the size of the muscle or the absolute tension developed. Later studies provided indirect evidence for a centrally-generated pressor response (McCloskey, 1981a; Leonard et al., 1985; Gandevia et al., 1993). In these studies, greater than expected increases in heart rate and blood pressure were observed when subjects attempted to contract muscles that were completely or partially paralysed by neuromuscular blockade. In each case, the cardiovascular response was closely related to the intensity, or effort, of contraction.

Perhaps the most compelling evidence for a centrally-generated pressor response is based on the experiments by Goodwin et al. (1972). Using tendon vibration to alter the excitability of the motoneuron pool artificially, they showed that for the same absolute muscle tension a contraction requiring a greater motor command produced a larger pressor response whereas a less demanding contraction resulted in a relatively weaker pressor response. Importantly, tendon vibration by itself does not produce a cardiovascular response (McCloskey et al., 1972).

It appears that the central motor command exerts its primary influence on the parasympathetic nervous system to produce a centrally-generated pressor response. This is based on observations that the characteristic increase in heart rate during attempts to contract pharmacologically-paralysed muscles voluntarily is blocked by atropine but not propranolol (Freyschuss, 1970; Victor et al., 1989), indicating that a central command signal results in inhibition, or withdrawal, of cardiac vagal tone. On the other hand, there is little or no increase in muscle sympathetic activity during attempts to contract paralysed muscles (Victor et al., 1989).

It is also proposed that along with the withdrawal of cardiac vagal tone, the central motor command simultaneously resets the arterial baroreceptors to a higher operating point (Gallagher et al., 2001; Querry et al., 2001; Ogoh et al., 2002). This adjustment would be an essential part of the exercise pressor response permitting sustained increases in heart rate and arterial blood pressure. A resetting of the operating point also allows the baroreceptors not only to regulate but also to prevent excessive increases in blood INTRODUCTION: PERFUSION 24

pressure during exercise. Incidentally, the resetting of the baroreceptor reflex’s operating point creates, momentarily, an apparent fall in arterial blood pressure. It is possible then that soon after the onset of exercise the central command achieves its effects on the circulation through the baroreceptor reflex, at least until the apparent fall in arterial blood pressure is reconciled (Rowell, 1993). Overall, it seems that the central command plays a key role in setting the pattern of cardiovascular activity at the onset of muscular exercise.

Figure 1.3 Central and peripheral contributions to the exercise pressor response During voluntary muscle contractions, a corollary signal that arises from and is proportional to the cortical motor command is sent to the cardiovascular control centres in the brainstem. The pressor response evoked by the corollary discharges is mediated primarily by an increase in heart rate. In addition, groups III and IV muscle afferents, which are sensitive to mechanical and chemical stimuli, contribute to the exercise induced pressor response by reflexively exciting the cardiovascular centres to increase heart rate and vasoconstriction. The evoked pressor response is monitored by an active arterial baroreceptor reflex.

INTRODUCTION: PERFUSION 25

Peripheral mechanisms Conclusive evidence that afferent inputs from active muscles reflexively excite the cardiovascular system was presented in the classic work of Alam and Smirk (1937). By employing local circulatory occlusion of the active muscles, they demonstrated that the exercise induced increase in arterial blood pressure persisted in the absence of further exercise for as long as the circulation remained arrested. They reasoned that the accumulation of metabolites in the contracting muscles was responsible for a reflex that raised blood pressure.

Later studies by Coote et al. (1971) and McCloskey and Mitchell (1972) provided a firm neuroanatomical basis for a muscle pressor reflex. First, both studies showed that electrical stimulation of ventral roots in cats, thus bypassing the central motor command, produced a small increase in heart rate and a large increase in blood pressure. Second, these responses were abolished after cutting the corresponding dorsal roots. In their study, McCloskey and Mitchell (1972) were able to clarify which types of afferents were involved in the muscle pressor reflex. Anodal blockade of group I and group II afferents, responsible for conveying information from muscle spindles and Golgi tendon organs, did not prevent the contraction induced increase in cardiovascular function. However, preferential blockade of groups III and IV afferents with local anaesthesia did abolish the reflex cardiovascular response to muscle contraction.

Both groups III and IV afferents, which include nociceptor afferents, respond to muscle contractions. Generally, group III afferents have been assessed to be sensitive to mechanical stimuli (Paintal, 1960; Kaufman et al., 1983) whereas group IV afferents are deemed to be sensitive to chemical stimuli (Kaufman et al., 1983; Mense & Stahnke, 1983). There also appears to be populations of groups III and IV afferents that are sensitive to both chemical and mechanical stimuli (Kniffki et al., 1981). Consistent with these properties, group III afferents are only transiently active at the onset of static contractions, whereas group IV afferents respond with latencies of 5 – 30 seconds, depending on the magnitude of contraction, and remain active throughout the contraction (Kaufman et al., 1983).

Chemosensitive and mechanosensitive afferents both project to the cardiovascular control centres in the brainstem (Iwamoto et al., 1985) but seem to engage different INTRODUCTION: PERFUSION 26

components of the autonomic nervous system. Chemosensitive afferents are favoured to activate sympathetically-mediated vasoconstriction reflexively provoking an increase in arterial blood pressure. This view is based largely on the observations that muscle sympathetic nervous activity (MSNA) has a delayed onset during static contractions, ranging from 1 to 2 minutes (Mark et al., 1985; Saito et al., 1986; Seals, 1989b), and remains elevated during post-exercise circulatory occlusion (Mark et al., 1985; Wallin et al., 1989). The delayed onset of MSNA is presumably due to the time required for metabolites to accumulate sufficiently to activate the chemosensitive muscle reflex. Furthermore, the rise in blood pressure and MSNA during contraction and post-exercise muscle ischaemia seems to be related to the size of the muscle mass exercised (McCloskey & Streatfeild, 1975; Freund et al., 1978; Seals et al., 1983; Seals, 1989a).

The discharge properties of mechanosensitive muscle afferents, which respond rapidly but transiently at the onset of muscle activity, would argue against their role in sympathetically-mediated vasoconstriction. Instead, evidence from studies involving electrical stimulation of the muscle or passive muscle stretch supports a role in cardiac acceleration. Several studies have shown that electrical stimulation of the muscle, to bypass the central motor command, still results in an abrupt increase in heart rate that is similar to a voluntary contraction (Hollander & Bouman, 1975; Bull et al., 1989; Friedman et al., 1992) and cannot be explained by a slow metabolic stimulus. Likewise, passive stretch of the muscle at intensities that do not result in a build up of metabolites produces a rapid increase in heart rate (Stebbins et al., 1988) that can occur independently of a change in blood pressure (Gladwell & Coote, 2002). These reflex cardiac responses are presumably achieved through the withdrawal of cardiac vagal tone, since the response to electrical stimulation is blocked by atropine (Hollander & Bouman, 1975) and the variability of successive heart beat intervals, an index of cardiac vagal activity, is reduced during passive muscle stretch (Gladwell & Coote, 2002).

These findings on the behaviour of groups III and IV muscle afferents indicate a differential control of the cardiovascular system during exercise. Mechanosensitive muscle afferents are involved in the regulation of cardiovascular activity at the onset of muscle contraction in parallel with the central motor command, whereas chemosensitive muscle afferents play a significant role in maintaining cardiovascular function during muscle contractions. Moreover, chemosensitive and mechanosensitive muscle afferents INTRODUCTION: PERFUSION 27

appear to produce independent cardiovascular responses. This was demonstrated by Fisher et al. (2005) who showed that the cardiovascular response to passive stretch of the calf muscles was not affected by concurrent post-exercise circulatory occlusion at different intensities of muscle contraction.

Muscle perfusion and muscle performance Much research has focussed on the reflex cardiovascular consequences of muscle activity; that is, the increase in heart rate and arterial blood pressure that support an increase in muscle work output. Several studies have investigated the other side of the story; the effects that changes in muscle perfusion pressure and blood flow have on muscle performance. Hobbs and McCloskey (1987) demonstrated in the isolated cat soleus muscle that stimulated force output was sensitive to changes in muscle perfusion. Force output declined in accordance with a fall in muscle perfusion. In the human adductor pollicis muscle, stimulated tetanic force output showed a similar dependence on perfusion pressure within the physiological range (Fitzpatrick et al., 1996b), indicating that local autoregulatory mechanisms, believed to operate across the physiological range of blood pressures, do not restore perfusion to a level that maintains muscle performance. During a sub-maximal voluntary contraction at a steady force output, this muscle fatigues faster and central blood pressure increases more through the greater central pressor response (Wright et al., 1999). This increase in central blood pressure offsets approximately half of the decline in force output (Wright et al., 2000).

The mechanism responsible for a perfusion dependent change in muscle force output could potentially involve a wide range of factors. Hobbs and McCloskey (1987) suggested that this effect could be explained by the alteration in oxygen delivery to the muscle. Contraction induced changes in the concentration of metabolites, interstitial ions, pH, as well as oxygen itself could contribute to the level of muscle oxygenation through their effects on the regulation of blood flow. Moreover, the generation of adenosine triphosphate for the energy-dependent excitation-contraction chain is directly affected by the availability of oxygen.

Muscle perfusion and standing The effect of muscle perfusion on force output demonstrated in cat soleus muscle and a human hand muscle might have an important role in human standing. Upright balance INTRODUCTION: PERFUSION 28

relies on soleus muscle activity that is subject to very large changes in perfusion pressure. Indeed, Egana and Green (2005) demonstrated an improvement in the endurance of the calf muscles with increasing angles of head-up tilt. However, the direct relationship between muscle perfusion and force output in human leg muscles has not been determined. Extrapolating this behaviour from a small hand muscle is not quite so straightforward since the leg represents a much larger muscle mass and its vascular beds are frequently exposed to higher blood volumes, blood flow and blood pressures. These inherent differences may modify how lower limb muscles respond to changes in perfusion pressure, particularly during standing where the soleus muscle must work for prolonged periods under a large head of pressure.

The regulation of the cardiovascular system during standing is normally attributed to the compensatory mechanisms that respond to the fall in arterial blood pressure that occurs when assuming an upright posture. This predominantly involves vasoconstriction of the arterial circulation to preserve mean arterial blood pressure (Tyden, 1977). In this case, vasoconstriction would directly oppose any increase in muscle perfusion. However, often overlooked in discussions of human postural control is that standing represents muscular exercise, which in itself, is capable of producing robust changes in the cardiovascular system to cope with the anticipated increase in muscle activity. The role of the exercise pressor response and its interactions with autoregulatory or compensatory mechanisms for perfusion pressure in human standing are not understood. Conclusions

The physics of human standing and the ease in which we perform this task belie the complexity of the physiological processes involved. The exact nature of the neural signals that control human standing have yet to be identified; however, the recent discovery of paradoxical muscle behaviour suggests that undiscovered central mechanisms are important. Postural control may be related to the cortical processes that generate muscle contractions and the accompanying perception of muscle force. Like all motor actions, standing must rely on the integration of sensorimotor and cardiovascular control processes to keep the body balanced. This could occur within the central nervous system as it is clear that corollaries of the central motor command guides muscle contraction and

29

regulates cardiovascular activity, although there is overwhelming evidence that shows that peripheral signals are equally important in maintaining muscle performance. Research plan

This thesis explores two areas of physiology that has traditionally been considered separately: motor control and cardiovascular control during human standing. Common to both of these physiological systems is the contribution from cortical control processes that are associated with the descending motor command. Psychophysical and physiological experimental approaches will be used to explore these broader aspects of the control of human standing.

Perceptions of force will be determined by lifting weights with the thumb, in line with current literature, to resolve the contributions of central and peripheral mechanisms (Chapter 2). The role of the central motor command in standing will be determined by a psychophysical assessment of force applied by the leg muscles during standing and matched non-standing tasks. These results are related to coherent activity between the motor cortex and the leg muscles during standing (Chapter 3). The significance of these results is examined by determining the effects of changes in blood pressure on the contractile performance of the leg muscles (Chapter 4) and to changes in blood pressure that occur during standing and performing equivalent non-standing contractions of the leg muscles (Chapter 5). 30

Chapter 2

On the sense of force and weight

Signals generated within the brain contribute to our perception of the external world. This is considered to be particularly significant for the perception of heaviness and muscular force exerted. For example, as a muscle is weakened by fatigue or partial paralysis, the increase in the motor command needed to lift a weight is thought to explain its increasing subjective heaviness. Contrary to this view, I show here that peripheral signals normally underlie this sense of exerted force, although it is possible to use a central signal if peripheral signals are unavailable. With different fatiguing contractions, the maximum force output of the thumb flexor muscles was halved and they were then used to lift an object. For healthy subjects this resulted in objects feeling the same or lighter, consistent with the expected effects of the conditioning contractions on the sensitivity of peripheral receptors. In contrast, for two deafferented subjects the perceived heaviness of the lifted object approximately doubled, in keeping with the central-signal theory. In separate experiments, the forearm muscles were completely paralysed with curare and allowed to recover to half-force output. This resulted in objects feeling lighter when lifted by the semi-paralysed thumb, even though the motor command to the motoneurons must have been greater. This is most readily explained by reduced lift-related reafference caused by paralysis of muscle spindle intrafusal fibres. It is concluded that peripheral signals, including a major contribution from muscle spindles, give rise to the sense of exerted force. In concept, however, reafference from peripheral receptors can also be considered a centrally-generated signal. These results therefore challenge the distinction between central and peripheral- based perception, and also the notion that muscle spindles contribute only to perception of limb position and movement. The concept of a corollary discharge, commonly used to explain force perception through pathways or mechanisms entirely within the central nervous system, can equally apply if we extend the concept to involve neural pathways traversing the peripheral nervous system to provide reafference.

SENSE OF FORCE 31

Introduction

Two sources of neural information could give rise to the sense of heaviness when we lift an object or force exerted by our muscles. Sensory receptors in the periphery respond directly to applied forces and a signal of effort is generated centrally as a corollary discharge of the motor command sent to the muscles. It is generally considered that the central signal provides the sensation of heaviness or weight when we lift an object (McCloskey, 1981b; Jones, 1986). This view comes from studies showing that when muscles are weakened by fatigue or partial curarisation, lifted objects feel heavier (McCloskey et al., 1974; Gandevia & McCloskey, 1976b, 1977a, b; Jones & Hunter, 1983a; Deeb, 1999). The increase in motoneuron drive required to generate the same force from the weakened muscle results in a corollary discharge signal that makes the same weight feel heavier (Gandevia & McCloskey, 1977a, b). Further evidence that central drive provides for this sense comes from the observation that neural drive to functionally related muscles increases the perceived heaviness of a lifted weight (Kilbreath & Gandevia, 1991).

There is a problem with the central command hypothesis as the exclusive determinant of perceived heaviness. Simply put, a muscle rendered half as strong does not make a lifted object feel twice as heavy, a concern expressed by McCloskey (1981b), a keen advocate of the corollary discharge hypothesis. Figure 2.1A shows the degree of muscle weakness along with the reported heaviness and the predicted from the curarisation study of Gandevia and McCloskey (1977b). Of course this discrepancy could be explained if the central signal is highly non-linear1 but the alternative hypothesis, which I will examine here, is that peripheral sensory receptors provide the dominant signal for perception.

Anaesthesia of the thumb makes a weight lifted by it feel heavier (Gandevia & McCloskey, 1977a; Gandevia et al., 1980), indicating that contact forces detected by

 1 Throughout this system, many small non-linear behaviours can be identified in isolation. However, complex, multi-element systems tend towards linearity in their overall behaviour. For example, the relationship between muscle force and EMG, which is considered to reflect the central motor command, is largely linear (Bigland & Lippold, 1954). Likewise, the postural reflex system has been shown not to deviate significantly from linearity although responses of individual afferent and motor units can be highly non-linear (Fitzpatrick et al., 1996). SENSE OF FORCE 32

cutaneous and perhaps joint afferents contribute to this sense. Golgi tendon organs, with their discharge that reflects the tension in muscles and tendons, should be a natural source for information about the heaviness of a load. Specific evidence that tendon organ afferents project to the sensory areas in the cerebral cortex was provided by McIntyre et al. (1984), indicating that these signals are likely to be perceived consciously. However, the contribution of intramuscular receptors has been difficult to determine as blocking their sensory signals necessarily blocks muscle contraction.

 Figure 2.1 Central and peripheral force signals A. Redrawn from Gandevia and McCloskey (1977b). The forearm was partially curarised to reduce handgrip strength. Maximum contraction strengths are shown in the lower graph. Above are weights selected with the normal arm to match 200 g lifted with the weakened arm. The orange “predicted” curve is an added calculation of the weights expected if judgement was based on a linear central command. B. Data on the left are tracings from Gregory and Proske (1979) of responses of cat Golgi tendon organ to a high-force (red) and low-force (black) stimulated muscle contraction (above), showing proportionally greater desensitisation with the high-force contraction. On the right, are tracings from Thompson et al. (1990) of the discharge responses of a unit to a muscle stimulated to contact at 5 N (above) after high-force (red) and low-force (black) conditioning contractions, showing a diminished response after high-force conditioning.

Tendon organ sensitivity and firing rates decline with a constant imposed tension (Gregory & Proske, 1975) and this can explain errors in force matching after brief forceful contractions (Thompson et al., 1990). This adaptation, illustrated in Figure 2.1B, varies with the level of tension developed with higher tensions producing proportionally larger declines in firing rates (Gregory & Proske, 1979). Thus, a high- force contraction over time should render tendon organs relatively unresponsive SENSE OF FORCE 33

compared with a low-force contraction. Similar adaptation also occurs in the discharge of slowly-adapting cutaneous afferents (Iggo & Muir, 1969) and probably muscle spindle primary afferents during intrafusal contraction (Cheney & Preston, 1976), although this is a complex matter for spindles as they have different forms of static and dynamic fusimotor drive.

Curare and its analogues are competitive antagonists at the neuromuscular junction and block muscle excitation. Thus, a central motor command can be issued but force output from the muscle is reduced. Sensory receptors and afferent neurons are unaffected and so will reliably transmit the true signal of reduced force output. The only exception to this comes from the neuromuscular junction block of muscle spindle intrafusal fibres as they share the same acetylcholine receptor. This indirectly affects the firing of spindle primary and secondary sensory endings during muscle activation in a time- and dose-dependent manner.

Beginning with these observations of the differential effects of muscle fatigue and curarisation on muscle contractile performance and afferent sensitivity, these experiments explore the origin of the senses of force, heaviness and effort with a series of crossed weight-matching experiments. In particular we test the current hypothesis that these senses are largely based on centrally generated signals. In normal subjects, fatiguing contractions of different force are made to affect peripheral receptors differently but create the same reduction in maximal force output. Two subjects with large-fibre peripheral sensory deafferentation are studied in this same way to determine how heaviness is perceived when peripheral signals are unavailable and only centrally- generated signals can contribute. In normal subjects, perceptions of heaviness are determined after partial recovery from prolonged total curarisation that paralysed both the extrafusal force-generating muscle and the intrafusal muscle spindle fibres but preserved other afferent sources. Methods

The experiments described here involve contralateral weight-matching tasks in which subjects used the unaffected indicator limb to match a weight lifted by the reference limb, before and after conditioning by fatigue, vibration or paralysis with curare. Another experiment involves a contralateral isometric force-matching task in which SENSE OF FORCE 34

subjects intermittently matched with the indicator limb the perceived force exerted by the reference limb in a sustained fatiguing contraction.

Subjects A total of sixteen healthy adults aged 22 to 59 years (mean 31.7) with no history of neurological disorder, and two with peripheral sensory loss, participated as subjects. Subject IW has a large-fibre sensory neuropathy (Sterman et al., 1980) with absent cutaneous and proprioceptive sensibility below the neck (Cole & Sedgwick, 1992). Since onset 26 years previously, he has regained good motor control but relies almost entirely on vision (Cole, 1995). Subject K has a large sensory-fibre peripheral neuropathy of 6 years affecting the limbs with a relative sparing of the trunk. He shows the coarse movement control of sensory apraxia although manages well when only a single limb is used. He has lost limb movement sense and position sense for the hands and wrists with better function at the elbows and shoulders. In the upper limbs, light cutaneous touch and vibration sense are absent with some sharp touch preserved.

Setup To undertake the contralateral weight-matching task, subjects sat with both forearms resting on a table and stabilised the hands by loosely wrapping the fingers around metal rods fixed to the table (Figure 2.2A). The thumbs were free to press down on a low- friction lever with no mechanical advantage and lift weights suspended out of sight below the table. Adjustable stops set an appropriate lever height for each subject and a movement range of 2 cm. To measure maximal contraction force, the weight was disconnected and a load cell was attached between the lever and the floor. This arrangement with the isometric load cells was also used for an experiment in which subjects matched isometric force during fatigue of one arm.

Design The design of these studies was to condition the lifting muscle in different ways. Fatiguing high-force and low-force contractions were made to identify the main effect of conditioning the muscle and afferents differently (Figure 2.2B&C). Further studies investigated the peripheral contributions to the changes in perceived heaviness that had been observed. Vibration applied to a contracting, but not passive, muscle causes Golgi tendon organs to discharge in proportion to the level of contraction (Brown et al., 1967; SENSE OF FORCE 35

Fallon & Macefield, 2007). Thus, in one experiment vibration was applied to the muscle during a low-force conditioning contraction to increase discharge rates with the aim of desensitising tendon organs further (Figure 2.2D). Control experiments were performed to examine the effects of vibration alone without fatigue. Finally, paralysis by curare was used to leave sensory receptors unaffected, other than muscle spindles, and examine perception when the same contraction force required a greater motor command (Figure 2.2E).



Figure 2.2 Setup and design A. Subjects flexed the thumbs to press down on a lever and lift weights suspended out of sight. A constant 500 g lifted by the reference hand was matched by adjusting the weight lifted by the indicator hand. Weight-matching tests were performed before and after conditioning the lifting muscle of the reference limb, with minimal delay between testing and conditioning. Five conditioning protocols were tested: B high-force fatigue in which the subject made a maximum isometric contraction until force had fallen to 40% of the initial maximum, C low-force fatigue with the 40% of maximal contraction (MVC) maintained until it could not be sustained, D the low-force fatiguing contraction was maintained while vibration was applied to the tendon of flexor pollicis longus, and E after a prolonged period of complete paralysis with curare and then allowing the muscle to recover maximal force output to 40% of the initial level.

Protocol Subjects simultaneously lifted a 500 g reference weight with one thumb and an indicator weight with the other. While lifting, the eyes were shut to exclude any visual SENSE OF FORCE 36

cues that might signal a weight difference. The reference side was randomised between subjects and kept throughout the experiment. Initial indicator weights varied randomly between half or double the reference weight, differences that were clearly detected on the first lift. After each lift, if they nominated that the indicator weight felt lighter or heavier than the reference, the weight was adjusted according to the subject’s instructions to increase or decrease. The increment or decrement was varied randomly by the experimenter (minimum 25 g) and could overshoot the target reference weight. When the weights were judged equal, the indicator weight was recorded as the subject’s perception of the reference weight. Twenty control matches were made prior to each experimental intervention, fatigue or paralysis, to establish baseline.

Fatigue. The experimental fatigue intervention was then carried out on the reference hand. Two methods were used to fatigue the thumb flexors: a high-force fatigue, in which the subjects maintained a maximal voluntary contraction (MVC), and a low-force fatigue in which subjects maintained a steady sub-maximal force at 40% of their maximum. Both were isometric contractions against a load cell that was connected between the lever and the floor. MVC force had been measured on a previous day as the highest of three brief maximal contractions made under verbal encouragement.

Both fatigue protocols ended when force output had decreased to 40% of the subject’s maximum. For the high-force fatigue, subjects were vigorously encouraged to maintain maximum force until it had fallen to the target 40% (Figure 2.2B). For the low-force fatigue, the steady 40% maximum force was maintained by visual feedback and stopped when it fell below the target and could not be restored despite verbal encouragement (Figure 2.2C). In another trial, low-force fatigue was accompanied by intense focal vibration over the flexor pollicis longus muscle (~2 mm at 50 Hz) using a tattoo machine with an 8 mm spherical bead (Figure 2.2D). Muscle vibration at this frequency and amplitude is near optimal to drive Golgi tendon organ responses (Fallon & Macefield, 2007). Control trials were made with vibration applied without fatigue, either with the muscle relaxed or making a 5% maximum contraction for the same duration as the low-force fatigue trial. Vibration of the muscle ceased as subjects began to match the reference weight. SENSE OF FORCE 37

At the fatigue end point, subjects immediately began a series of twenty weight matches with the same protocol as the pre-fatigue matches. After completion, maximum contraction force was measured again to determine the extent of recovery. The different testing protocols each lasted approximately 45 minutes and were performed at least two days apart.

The deafferented subjects could only be studied on a single occasion. Subject IW performed the high-force fatigue protocol and K performed the low-force protocol. The degree of fatigue was reduced so that the target was 50% maximum rather than the 40% target of the normal subjects. Both lifted with the eyes open, necessary to know that they had lifted the weight off the rest and not hit the stop. Subject K required the pads of the thumbs to be stuck to the levers with double-sided tape to keep the thumbs in position when he relaxed between lifts. The maximum contraction forces used to determine the target levels were performed before testing (4 hours for IW, 1 hour for K). Testing these subjects took longer and so only eight matches could be made in the period during which the normal subjects made twenty.

For the normal but not the deafferented subjects, force-matching experiments were undertaken to measure perceptions of force exerted during the course of the isometric fatiguing contractions. The subject made a brief isometric contraction with the indicator thumb every 30 s to match the perception of the force exerted with the reference thumb, which was making the high-force or low-force fatiguing contraction.

Paralysis. A non-depolarizing neuromuscular blocker, rocuronium, was used to paralyse all muscles distal to the cuff (Wierda et al., 1994). The rocuronium (5 – 7 mg; -1 0.1 mg.kg ) was in 50 ml of Plasma-Lyte (Baxter) with 2 ml of 8.4% NaHCO3 solution to balance pH. After the control experiments, an intravenous cannula was inserted proximal to the wrist. Blood was removed from the arm vasculature by elevation and compression by a rubber toroid rolled proximally before a dual-chamber cuff around the upper arm was inflated rapidly to 300 mmHg by regulated pressurised air. A retrograde infusion of the rocuronium solution over 60 s began within 30 s. Paralysis was complete within 10 minutes, verified by no detectable flexor or extensor force or EMG in the forearm muscles. The cuff remained on for 13 – 18 minutes and subjects generally developed some transient dizziness and diplopia on deflating the cuff. These symptoms SENSE OF FORCE 38

were short-lived (5 – 10 minutes) and the subject rested with the arm in a sling until partial recovery to 40% of initial maximum force. This recovery took 1 – 2 hours, during which brief maximal efforts of thumb flexion were made. When the target 40% force recovery was reached, subjects made twenty weight matches (Figure 2.2E). After completion, maximal contraction force was measured again.

Measurement and analysis Matches were not synchronised and subjects had idiosyncratic rates of matching. Thus, for purposes of averaging across subjects, matched weights of all subjects were pooled and responses divided into 50 s bins (except for the paralysis study that was binned at 100 s) backwards from the end of the pre-fatigue matches and forward from the start of the post-fatigue matches. The differences between the responses of the normal and deafferented subjects and force matching during and after fatigue are described quantitatively. Physiologically important effects of conditioning are apparent by visual inspection but ANOVA by epoch (before and after conditioning) and post-hoc Student’s t tests of bin data were applied to confirm statistical significance at P < 0.05. Mean data, generally as percentage changes after conditioning, are presented with 95% confidence intervals in the format: mean [low 95% limit, high 95% limit] (Curran- Everett & Benos, 2007). Results

Fatigue and heaviness Subjects matched the 500 g reference weight with reasonable accuracy (coefficient of variation 0.088) that did not vary with repeated tests (Figure 2.3A, left). Immediately after the high-force fatiguing contraction of the muscle lifting the reference weight, the mean indicated weight was 15.2% [6.8, 23.6] less. That is, the weight lifted by the fatigued muscle felt lighter (Figure 2.3A, right). With the muscle fatigued to 40% of initial maximal force subjects could still readily lift the 500 g reference weight. Over the following 500 s, the indicated weight recovered towards the pre-fatigued levels. In contrast, following low-force fatigue to the same level of 40% initial maximal force, there was no difference in the reported weight immediately afterwards. It was noted that there was a small but significant upward drift in matched weights over the 500 s. When vibration was applied to the tendon during the low-force fatiguing contraction, the mean SENSE OF FORCE 39

indicated weight immediately afterwards was 14.0% [3, 25] less. Vibration applied with the muscle relaxed or making a non-fatiguing weak 5% maximal contraction had no effect on perceived heaviness (Figure 2.3B). At the end of the post-fatigue matches, maximal contraction forces were more than 90% of initial levels for all subjects.

Figure 2.3 Effects of muscle fatigue on weight perception for normal subjects A. Weights selected to match 500 g lifted with the reference thumb before (left) and after (right) fatiguing the reference muscle until its force was reduced to 40% of initial maximum force. Data are mean ± S.E.M. for 9 subjects. After high-force fatigue (filled red) subjects indicated that the reference weight felt lighter and this recovered over several minutes. After low-force fatigue (open red) to the same level (40% maximal), they indicated that the weight felt the same as that lifted by the non-fatigued arm. After the same low-force fatigue but with strong vibration applied (open blue), they indicated that the reference weight felt lighter as it had with the high-force fatigue. B. Tendon vibration with the muscle at rest and during a weak sustained contraction (5% MVC) had no effect on perceived heaviness. * P < 0.05 from baseline by ANOVA.

An entirely different picture is seen in the subjects with large-diameter afferent sensory loss. Before fatigue, matching accuracy was not significantly different from the performance of the normal subjects (coefficient of variation 0.102; Figure 2.4, left). When the muscle was fatigued to 50% of initial force output, the lifted weight was initially reported as approximately twice as heavy (112% increase for IW with high- force fatigue; 92% increase for K with low-force fatigue; 102% increase combined) with recovery towards baseline occurring over approximately 10 minutes (Figure 2.4, SENSE OF FORCE 40

right), longer than the recovery in the other direction for the normal subjects.

 Figure 2.4 Effects of muscle fatigue on weight perception for two deafferented subjects Weights selected to match 500 g lifted with the reference thumb before (left) and after (right) fatiguing the reference side until its force was reduced to 50% of the initial maximal force level. The filled circles are subject IW who made a high-force contraction to fatigue the thumb flexors until force output fell to half the initial maximum. The open circles are subject K who fatigued by a sustained 50% maximal contraction until he could not hold it. For these subjects, halving strength caused the reference weight to feel approximately twice as heavy.

Fatigue and force exerted Figure 2.5 shows forces made by isometric contraction of the indicator thumb to match the perceived force exerted by the reference thumb during the fatiguing contractions. For the sustained maximal high-force contraction, the indicated perceived force declined immediately, following the actual decline in force output. However, indicated force was always greater than perceived force, declining at only 50 – 60% of the actual decline in force output (Figure 2.5, left panels). During the sustained low-force isometric contraction, the perception of force exerted rose steadily, reaching 53% [23, 90] higher than actual at the point of 40% maximal force output (Figure 2.5, right panels). At that time, perceived force fell, reflecting actual force exerted. SENSE OF FORCE 41

 Figure 2.5 Perceived force exerted during an isometric fatiguing contraction A. Data for one subject. On the left, the black trace shows the declining force of a continued maximal isometric contraction of the reference thumb. Blue are intermittent isometric contractions made with the indicator thumb to match the reference force. On the right, the reference contraction is at 40% of the pre-test maximal force until it cannot be held. At that time (break), force declines. The forces indicated by the contralateral indicator thumb increase during the period of steady force output but decline steeply when force output is maximal. B. In the format of A are the group mean (± S.E.M.) of the actual and indicated force for each fatigue protocol. To allow for different times to reach the prescribed fatigue level, time is scaled to the break point with relative equal-width bins averaged across subjects. Orange lines are the predicted perceived force exerted if perception is based on a central command with unity gain i.e. halving maximal force output by muscle fatigue requires doubling the central command to maintain force output, as indicated by the results of subjects IW and K in Figure 2.4.

Paralysis with curare Curare does not affect Golgi tendon organs or cutaneous receptors. Thus, any muscle weakness should bias perception to make lifted objects feel heavier if it is influenced by an increased central command or corollary discharge. The opposite observation was made. After complete and prolonged paralysis with rocuronium and recovery of maximal force output to 40% of the pre-paralysis maximum, subjects chose lighter weights with the unaffected indicator thumb to match the 500 g lifted with the weakened reference thumb (Figure 2.6A). The first post-paralysis match with the unaffected thumb was 31% [17, 45] less than the 500 g reference weight lifted with the thumb that had been weakened by 60%. Recovery over time brought the matched weights back towards the reference but after 500 s a residual shortfall remained. SENSE OF FORCE 42

Individual match data for one subject who made fast weight matching decisions shows a steady recovery of perceived force that parallels the muscle recovery (Figure 2.6B).

 Figure 2.6 Effects of paralysis with curare on weight perception A. Weights selected to match 500 g lifted by the reference hand, before and after paralysis with rocuronium. During the period “curare” the reference arm was completely paralysed and allowed to recover to 40% of initial maximal force before testing. Afterwards, lighter weights were chosen to indicate the heaviness of the weight lifted by the paralysed reference hand. Individual responses were binned into 100 s epochs before averaging. B. Individual matches for one subject superimposed on the maximal contraction force before and after paralysis (red bars relative to the initial force on arbitrary scale). The reduction in matched weight is proportionally less than the reduction in maximal contraction force but recovery follows the force recovery. * P < 0.05 from baseline by ANOVA.

Discussion

Von Helmholtz (1867) proposed that sensations of innervation are generated within the brain as part of the process of activating the muscles. It is generally accepted that this internal signal, or corollary discharge (Sperry, 1950), gives rise to the sensation of force exerted by muscular contraction (McCloskey, 1981b; Jones, 1986). Sherrington held a contrary view: “that during a willed movement the outgoing current of impulses from brain to muscle is accompanied by a sensation for innervation .... remains unproven” (Sherrington, 1900; see Matthews, 1982 for review). The heavier weights chosen by the deafferented subjects to match the weight lifted by the fatigued limb suggest that some form of central signal can be available. However, it is clear from the entirely different SENSE OF FORCE 43

matching behaviour of the normal subjects that we do not normally perceive force by that mechanism. The results show that with fatigue and paralysis objects can feel lighter, opposite to the predictions of perception based on a centrally-generated signal (McCloskey et al., 1974; Gandevia & McCloskey, 1976b, 1977a, b; Jones & Hunter, 1983a; Deeb, 1999). From these results I will argue that both the central and peripheral views are correct and that in large part the centrally-generated corollary discharge travels via the periphery to return as reafference from muscle receptors.

Here I propose a unifying hypothesis in which the weight-matching task is achieved using both efferent and afferent information. The hypothesis relies on the concept that with every volitional motor command there is an associated expectation of reafference arising from the consequences of the motor output. This expectation could come from a memory of past actions or from operation of an internal model. Deviations from the expected reafference may be interpreted as indicating that the load being acted on is not as predicted. With this, subjects are able to gauge whether the load acting against one limb is the same as that acting against the other. In the following we interpret our results in the light of this model and attempt to reconcile the findings of others using the same hypothesis. Figure 2.7 summarises the results and interpretations of the different studies.

Paralysis Roland and Ladegaard-Pedersen (1977) weakened the muscles of one arm by partial neuromuscular blockade using gallamine. When asked to match forces between limbs, subjects did not indicate that they applied more force with the weakened limb, although this would have generated a greater corollary signal of effort. This was taken to indicate that a peripheral signal, probably from Golgi tendon organs, gives rise to a sensation of force. When asked to match effort, however, subjects generated a force with the unaffected side inversely proportional to the residual strength of the weakened muscle, thus showing that the central effort signal gives rise to a distinctive sensation. At the same time, Gandevia and McCloskey (1977b) reported that weakening the lifting muscle by partial neuromuscular blockade with a natural curare compound (D- tubocurarine) makes objects lifted by that arm feel heavier. They concluded that this was evidence for the corollary or efference copy of the central motor command dominating the sense of heaviness of lifted objects. SENSE OF FORCE 44

 Figure 2.7 Interpretation: a model of force perception based on fusimotor reafference In all figures (A-D), the weight is lifted by equal amounts of muscle shortening. The key point is that the central command signal (Action) was increased by similar amounts in the deafferented (B), paralysis (C) and fatigue (D) experiments. The thickness of each pathway indicates the relative strength of the signal. A. Implicated in the perceptions of force and weight are central corollary discharge (cd) and/or efference copy (ec) signals, and peripheral signals largely from tendon organs (Ib). a-g linkage is indicated at supraspinal and spinal levels but the site is not critical for this interpretation. The results of these studies show that spindle afference (Ia & II) contribute to the perception of force. B. Deafferented subjects have a quality of force perception that is different to normal and based entirely on the central signals. These double in size after fatigue so that weights are reported twice as heavy compared with the reference side (ref). C. During the early phase of curarisation, as tested by Gandevia et al., (1977a), relative extrafusal paresis (left) sees a greater spindle receptor stretch through intrafusal shortening as the drive is linked to the increased drive required to offset extrafusal weakness. Increased spindle reafference creates a perception of greater force exerted relative to the reference side. During recovery from complete paralysis, relative intrafusal paresis results in spindle unloading so that reduced spindle reafference creates a perception of less force exerted. D. After high- force fatigue, greater receptor desensitisation reduces spindle and tendon organ reafference (left) so that weights feel lighter than the reference and after low-force fatigue (right). SENSE OF FORCE 45

How can these results be resolved with the observations here that objects feel lighter after paralysis? After recovery to 40% maximal force, the motor command sent to the muscle would more than double if an efference-copy signal was used to judge heaviness centrally, yet objects were perceived to be one-third lighter. Curare does not cross the blood-brain barrier and subjects reported no sensory or motor changes beyond the paralysed forearm so we can exclude a direct central effect. Some spill-over of curare on releasing the tourniquet might have created a transient weakness of the reporting hand. We expect that any significant weakness had resolved by the time of testing (> 1 hour) when recovery on the paralysed side had reached 40% maximal force. However, according to the central signal hypothesis, any weakness of the reporting hand should lead to lighter matches of objects lifted on the reference paralysed side and therefore the result we show here would underestimate the true effect.

Complete muscle paralysis by curare is a dynamic process that affects both extrafusal and intrafusal fibres. Block of intrafusal fibres requires higher doses of curare and the effect lags the block of extrafusal neuromuscular junctions (Carli et al., 1967; Smith & Albuquerque, 1967; Emonet-Dènand & Houk, 1968). The data of Smith and Albuquerque (1967), replotted in Figure 2.8, show that repeated ventral root stimulation during curare infusion results in an increasing discharge from muscle spindle primary afferents during and after the period when twitch force declines. Some time after zero force output has been reached, spindle discharge begins to decline and eventually does not respond to the contraction. The explanation is that the extrafusal fibres, being blocked first, do not unload the spindle during contraction and the effect of the intrafusal contraction on the primary afferents is revealed. After the extrafusal fibres have been blocked completely, the intrafusal block catches up and afferent discharge declines.

As with the onset of paralysis, the recovery of intrafusal fibres from paralysis by curare also lags that of the extrafusal fibres (Carli et al., 1967; Emonet-Dènand & Houk, 1968) consistent with the spindle capsule acting as a diffusion barrier (Matthews, 1972). Gallamine is a synthetic non-depolarising neuromuscular blocker with the triple quaternary amine structure of the curare compounds and a similar transport rate 0 constant (ke ) between plasma and effect (rocuronium 0.14; gallamine 0.16; D- tubocurarine 0.17) and requires similar effective plasma concentrations (Proost & SENSE OF FORCE 46

Wright, 2001). Yamamoto et al. (1994) demonstrate this delayed fusimotor paralysis and recovery with gallamine and propose that a fundamental receptor difference as well as a diffusion-lag across the spindle capsule (Matthews, 1972) are the underlying mechanisms. In addition, at lower doses, the pharmacokinetics of these competitive antagonists causes a rapid uptake and concentration at the neuromuscular junction and the brief paralysis is due to drug redistribution (Goodman & Gilman, 1996).

 Figure 2.8 Effects of curare on spindle discharge during ventral-root stimulation Results of an experiment by Smith and Albuquerque (1967) in which the ventral root of an isolated muscle was repeatedly stimulated with brief tetani to evoke isometric contractions. Curare injected at points “c” resulted in complete paralysis over 12 minutes, the force shown in the lower panel. Spindle IA afferent discharge rates were measured during contractions (stim: filled blue circles) and between contractions (rest: open circles). During the period of paralysis, discharge rates increased, reaching a peak after complete paralysis. The explanation is that extrafusal unloading of the spindle is progressively blocked, revealing the effect of intrafusal contraction on afferent discharge. The intrafusal paralysis is delayed and leads to a later decline in IA discharge until it does not rise above resting levels (point ). Beyond this point we can predict that extrafusal recovery ahead of intrafusal recovery will keep IA discharge low, perhaps less than resting levels, by unloading the spindles until intrafusal contraction recovers.

These pharmacokinetic differences between the extrafusal and intrafusal effects of curare mean that during induction, and with low doses, partial paralysis results in a loss of spindle unloading during contraction. The uncompensated -mediated intrafusal shortening increases primary (and probably secondary) spindle reafference (Figure 2.8, ascending limb). Thus, the subject receives an increasing signal from the spindles in response to the motor command that was issued. After complete extrafusal paralysis, the SENSE OF FORCE 47

delayed intrafusal paralysis results in a declining afferent response until eventually there is no response and spindle afferent firing rates during contraction do not rise above resting rates (Figure 2.8, descending limb). During recovery from paralysis, the pharmacokinetic time lag on the intrafusal fibres means that extrafusal shortening will exceed the intrafusal and tend to unload the spindles.

The most straightforward explanation for the finding that a lifted weight feels lighter after complete paralysis is that reafference from muscle spindles contribute significantly to the sense of heaviness and force exerted. Curare has paralysed the intrafusal muscle fibres so that the fusimotor co-activation that accompanies the extrafusal contraction creates only a small or no reafference from the spindle afferents, with a possibility that the volitional contraction is associated with a reduction in spindle firing. The results of Gandevia and McCloskey (1977b) are likely to reflect these pharmacodynamics of curare (Figure 2.1A). With transient partial paralysis, the extrafusal effect would dominate initially and a lagged effect on the intrafusal fibres could have resulted in the relatively rapid disappearance of the increased heaviness they observed. Figure 2.1A shows that by the time subjects’ perceptions of heaviness had returned to baseline the force output of the extrafusal fibres had only recovered by one- third. That is, in addition to the matched weights not reaching predicted values, they also returned to baseline too early, consistent with a delayed intrafusal paresis and reduction in the reafferent spindle signal.

Fatigue Numerous studies have shown that muscular fatigue induced by isometric contractions increases the perception of force exerted and the heaviness of lifted objects (McCloskey et al., 1974; Gandevia & McCloskey, 1976b, 1977a; Jones & Hunter, 1983a; Deeb, 1999). As muscle contractility decreases, motoneuronal drive must increase to maintain force output. The central corollary discharge hypothesis proposes that the greater descending corticospinal drive to support the contraction gives rise to the greater sensation of force exerted. In the absence of large fibre afferent inflow, lifted objects feel twice as heavy (Figure 2.3) when lifted with a fatigued muscle that can only produce half of its original force. In these subjects we can probably assume that we are seeing the pure central SENSE OF FORCE 48

corollary discharge sense2 in the absence of an afferent contribution and that this is the increased heaviness predicted by the central corollary discharge hypothesis. This gain suggests that the process, overall, is approximately unity. That is, half the strength requires double the motor command to generate the same force output. Despite uncertainties about the processes of central perception, central fatigue, and central adaptation during a sustained contraction, this gain means that we can hazard a guess at the sense generated by central corollary discharge during fatiguing contractions. A basic model indicates that perceived force should remain unchanged as muscle force output is lost during a sustained maximal effort, and should increase steadily to a maximum as fatigue reduces maximal force output during a sustained sub-maximal contraction at constant force (see Figure 2.5B). On the other hand, a basic model of sensation produced by afferent inflow alone should follow actual force output with an allowance for time-varying effects of sensor transduction. Using a constant-force protocol similar to that of the right panel in Figure 2.5B, McCloskey et al. (1974) showed that the perception of exerted force rose during the period in which constant force could be maintained. Although taken as evidence for perception through the central corollary discharge signal, their data also indicate a response less than unity and closer to the actual force exerted than the central gain determined in these studies. In the present study, perceptions of force were obtained beyond the point at which the target force could not be maintained. From this point onwards, perceived force fell dramatically and paralleled the fall in actual force output. This is inconsistent with the central-signal model and indicates a dominant input from afferent signals rather than a central signal.

The firing rates of Golgi tendon organs decline through desensitisation during sustained force and this effect is greater at high force levels (Vallbo, 1974a; Gregory & Proske, 1975, 1979). The same applies to muscle spindle primary afferents with applied stretch (Vallbo, 1974a). Following high-force fatigue to a level of 40% maximal force output, objects actually felt lighter whereas perceived heaviness was unchanged after the low-force fatigue. Tendon vibration during a low-force fatiguing contraction had an effect similar to high-force fatigue in that objects felt lighter afterwards. Vibration could  2 It is worth noting that subject IW in particular, who after many years has good insight to his condition (see Cole, 1995), did not describe his sense as one of heaviness. It had a different and less overt quality that he had learned to equate with heaviness and force. SENSE OF FORCE 49

have multiple and complex effects. A segmental tonic vibration reflex could mean that a smaller central command was required during the fatigue conditioning but greater desensitization of spindle receptors and the loss of afferent support for the contraction could require a greater central command. Vibration during contraction drives Golgi tendon organs (Brown et al., 1967; Fallon & Macefield, 2007) and could increase their desensitisation to resemble that of a high-force contraction. Thus, the different effects of high-force and low-force fatigue as well as the effect of vibration are again consistent with the dominance of afferent inflow for the sense of exerted force.

The afferent signal The results of the paralysis study specifically implicate reafference from muscle spindles as a major contributor to the sense of exerted force and the heaviness of lifted objects. The results of the fatigue study are also in keeping with a significant spindle contribution. With progressive fatigue, the increasing motor command to maintain force output results in greater fusimotor drive. However, through different processes that include receptor and mechanical adaptations, muscle spindles become less responsive to fusimotor drive so that the discharge frequency of spindle primary afferents declines progressively during a sustained isometric contraction, roughly halving after a minute during a 20% MVC contraction but more for higher initial firing rates (Macefield et al., 1991).

With firing rates largely proportional to muscle force, Golgi tendon organs appear to be the natural candidate for signalling exerted muscle force (Gregory & Proske, 1979). Cutaneous slowly-adapting receptors also provide a signal proportional to the contact force and contribute to the sense of heaviness of a lifted object (Gandevia & McCloskey, 1977a; Gandevia et al., 1980). Applying the concept and nomenclature of von Holst and Mittelstaedt’s (1950) reafference principle, each of these receptor systems will return a reafference related to the force of the muscular contraction but each has a different relationship with externally imposed perturbations. Exafference from muscle spindles largely reflects imposed changes in muscle length and its time derivatives, exafference from Golgi tendon organs largely reflects imposed forces along the axis of the active muscle, whereas cutaneous exafference reflects the location, strength and direction of imposed forces. With the different receptor types contributing to the different aspects of perception, it seems likely that the central nervous system SENSE OF FORCE 50

makes a compromise across the afferent and efferent signals related to the motor action to generate the sense of heaviness of a lifted object.

The studies here show a strong contribution of muscle spindle reafference driven by fusimotor action to the perception of force exerted. This suggests that the corollary discharge or efference-copy signal that is thought to give rise to these sensations may not be exclusive to the central nervous system but also can significantly involve the peripheral nervous system. Although shown here for the muscle spindle system, all receptor classes could provide this “reafference corollary discharge”. Thus, the apparently conflicting views of von Helmholtz (1867) and Sherrington (1900) can be resolved if we consider that the efference-copy or corollary-discharge pathways can include passage through peripheral as well as central pathways.

Conclusions The key conclusion of this study is that the senses of force exerted and the heaviness of lifted objects are largely derived from peripheral afferent inflow, and this takes the form of reafference related to the command sent to the muscles. Muscle spindles contribute significantly to this reafference although other sensory receptors responding to muscle contraction can also contribute. The next chapter examines the sense of exerted force during upright standing, where the weight lifted is one’s own body. 51

Chapter 3

The sensation of standing

With the perception of force being intimately linked to the peripheral consequences of intended motor actions and considering the strong continuous plantar torque that is required to stop the body toppling it is unusual that standing does not generate a greater sensation of muscle activity. Generally, we do not have a sense of the force we exert with the calf muscles. This phenomenon is investigated through a series of studies that first quantifies the perceived force during human standing and then seeks to determine the physiological basis for this behaviour. Sequential force-matching tasks in different postures reveal that passive force from stretch of elastic tissue at the position of standing provides about half of the total torque used to stand, and this is not perceived as exerted force. Of the actively generated force that produces the other half of the total force, subjects perceived the contraction in the calf muscles to be one-third that of the balancing torque during standing. This underestimation in force was not due to the compressive pressure of the body acting on the feet and nor was it influenced by changes in passive muscle tension. A comparison of corticomuscular coherence shows that the strength of coherence during standing was one-third that when making a voluntary contraction at the same force output. This indicates that the underestimation in force is related to the relatively small cortical activation of the muscle during standing. The presence of vestibular reflex activity only during standing indicates that this sub-cortical drive is only activated during standing and accounts for much of the balancing torque. This indicates that the altered perception of force during standing is due to the sub-cortical activation of the calf muscles to which the cortex does not have perceptual access. Without the efference copy of the sub-cortical drive available, the cortical perceptual processes cannot extract the reafference signal that feeds the sense of exerted force.

SENSATION OF STANDING 52

Introduction

With the entire weight of the body bearing down on them, extremely large forces act on the feet when we stand compared with sustained forces we apply in other activities. The upward thrust force is usually divided between the two legs and directed approximately at the centre of mass of the body when we stand still. Of course, we do not have to resist these forces by muscular contraction. Only the muscular contraction required to prevent the body from toppling is necessary and, as the centre of mass is almost always in front of the ankles, a sustained contraction of the calf muscles is required. The mean level of this contraction is not insignificant and amounts to 15 – 20% of the contractile force of the soleus muscles, the main muscle active at the ankles during quiet standing (Joseph & Nightingale, 1952; Smith, 1954; Portnoy & Morin, 1956). This muscle contraction results in a net torque applied to the ground with the front of the feet pressing down around the axis of the ankles in the same way that the thumbs pressed down against a resistance in the previous chapter. With this in mind, it seems unusual that we can stand for long periods and feel quite light on our feet. Standing is not perceived as particularly difficult or requiring an effortful or forceful contraction.

When we apply a force by pressing against a resistance or lifting an object, a centrally generated signal, the corollary discharge of the motor command sent to the muscles to perform the action, is associated with the sensation of force and effort. Although it has generally been accepted that this central signal rather than a peripheral afferent signal is the dominant means by which we judge force and heaviness, the results of the previous chapter indicate that the dominant source of information is indeed through peripheral afferents and that, through reafference, it actually represents the major corollary discharge signal. Whether this perceptual process of direct volitional actions also applies to more automatic postural and balance actions is not certain as the muscles appear to be influenced by different supra-spinal and spinal pathways in these tasks. For example, the soleus H-reflex, the electrical analogue of the short-latency stretch reflex, is reduced for standing when compared to the same electromyographic activity lying supine (Chalmers & Knutzen, 2002). There is also evidence that the H-reflex is inhibited presynaptically during standing (Katz et al., 1988; Koceja et al., 1993). Galvanic stimulation that produces a stereotyped vestibular evoked muscle reflex and postural response during standing (Coates, 1973; Nashner & Wolfson, 1974; Lund SENSATION OF STANDING 53

& Broberg, 1983; Britton et al., 1993) is absent when an equivalent muscle contraction does not involve balancing the body (Fitzpatrick et al., 1994a). Reports on corticospinal excitability during standing are mixed. Studies that have used transcranial magnetic stimulation to evoke muscle responses during standing indicate, perhaps paradoxically, an increased corticospinal excitability compared with similar volitional tasks (Ackermann et al., 1991; Obata et al., 2009; Tokuno et al., 2009).

The series of studies presented here examine the perceptions of force and effort during human standing. First, Study 1 shows that we only perceive a small fraction of the contractile force we exert to support the body. Further studies seek explanations for this phenomenon. Study 2 examines the hypothesis that the sensation is masked by other sensory input that originates from the large weight of the body bearing down through the feet. Study 3 examines the hypothesis that the sensation is reduced because a large level of the ankle torque used to stand is generated by passive elements rather than contractile action. Studies 4 and 5 examine the hypothesis that significant motor drive to the leg muscles during standing has a different, perhaps sub-cortical, origin not accessed by perceptual centres.

As five individual studies are presented here, the traditional Method and Results presentation would become overbearing and therefore this work is presented as a mini thesis, with the individual studies presented in their historical order, as hypothesis, method, result and conclusion, before a general discussion. These experiments were conducted over a couple of years on a total of 35 healthy subjects, aged between 19 and 60 years (mean 31.0). Study 1: The perceived force of standing

Subjects stood on a force platform with the eyes open and bare feet side-by-side separated by ~10 cm (Figure 3.1A). When settled, they were asked to attend to the torque they exerted against the floor (platform) produced by the contraction of the calf muscles. After 10 s, a rigid support was quickly positioned and subjects were strapped to it securely around the pelvis and chest (Figure 3.1B). This took 5 – 10 s. The support had been previously arranged for each subject so the body and ankle alignment was held in the same posture as during standing. Once supported, they relaxed the leg muscles completely, which were monitored by surface electromyography (EMG), and were then SENSATION OF STANDING 54

asked to produce a contraction of the calf muscles to push down against the floor, matching the force (torque) they had perceived when standing freely. No specific instruction was given to match either the effort or the tension developed within the muscles. Ten trials were made with rest periods between. Eight subjects were studied.



Figure 3.1 Perceived force during standing A. Subject freely standing on a force plate with electromyography (EMG) of the right soleus and tibialis anterior (TA), or B. supported upright by straps around the chest and waist. C. Raw data for a single trial. The subject was standing freely for 10 s before their body was supported upright (first vertical line), and when relaxed, matched the ankle torque perceived during standing (second vertical line). D. Group mean data (± S.E.M.) of the perceived torque during standing as a percentage of the exerted active torque. **P < 0.01 by ANOVA.

Ankle torque was recorded from a force platform (Kister 9286B, Winterthur, Switzerland) and EMG of the right soleus and tibialis anterior were sampled at 2 kHz. Passive torque was estimated as the residual torque recorded by the force platform when no EMG was present in the muscles. Active torque was calculated as the additional torque generated by active muscle contractions above the passive baseline. The maximum ankle plantar torque was determined for each subject in a seated position with both legs extended (posture shown in Figure 3.4A).

Data from a representative subject for a single trial are shown in Figure 3.1C. Passive ankle torque was 21.7 Nm, which represented 55.6% of the total torque of SENSATION OF STANDING 55

39.0 Nm. Actively generated torque in excess of the passive was 17.3 Nm. When matching the standing ankle torque, the subject made a much smaller contraction and generated just 5.3 Nm above the passive torque baseline or 30.7% of the active torque during standing. A similar behaviour was observed across subjects (Figure 3.1D). Passive ankle torque represented 49.6% [41.1, 58.1; 95% confidence interval] of the total ankle torque generated during standing. Subjects underestimated the torque generated during standing with active perceived torque just 31.7% [28.6, 34.8] of the active torque during standing (P < 0.01 by repeated-measures ANOVA).

A temporal effect? To control for a possible systematic temporal effect in this sequential force-matching protocol, ability to reproduce an isometric force after an elapsed time was determined. Subjects were supported upright throughout (i.e. never freely standing) and the initial contraction torque was a visual display target on an oscilloscope set to match the torque exerted during free standing (Figure 3.2). The time- course of the contract-relax-match during the standing trial was used and subjects reported perceived torque in the same way. During the relax-match period, the oscilloscope target was extinguished.



Figure 3.2 Effect of delay on reporting perceived force A. The subject was supported upright throughout each trial. B. Electromyographic (EMG) activity of the soleus muscle and the target ankle torque were presented to the subject to maintain the same torque exerted during standing. Visual targets were removed at the relax period (first vertical line). After the reporting delay, subjects matched the perceived force without the visual target (second vertical line). C. Group mean data (± S.E.M.) of the perceived active torque as a percentage of the exerted active torque. SENSATION OF STANDING 56

Subjects could accurately perceive and reproduce the force exerted by the plantar flexors when matched sequentially, although the repeated ankle torque tended to overestimate the initial by a mean of 10.1% [-0.1, 20.3].

Conclusions. Approximately half of the ankle torque during standing comes from mechanisms that generate passive stiffness. The torque generated by active contraction during standing is perceived as approximately one-third that of an equivalent volitional contraction. Study 2: Pressure under the feet?

This experiment tests the hypothesis that the pressure under the feet produced by the weight of the body influences, and perhaps masks, the perception of force exerted during standing. A contralateral weight-matching task was performed with and without a large normal reaction force under the reference leg. Eight subjects were studied.

Experiments were conducted with the subjects seated, hips and knees flexed at 90°, and each foot resting on a separate foot plate that rotated about the axis of the ankle. Hidden from the subject, weights were attached to the rear of each foot plate so that they were lifted by plantarflexion of the calf muscles (Figure 3.3A).

A variable test weight lifted by one foot was compared to a 2.0 kg reference weight lifted by the other, with the test and reference sides being randomised across subjects. The initial test weight varied from 50% to 200% of the reference weight. Subjects were instructed to lift the reference and test weights simultaneously and indicate whether the test weight felt the same, lighter or heavier than the reference weight. When the subject relaxed, the experimenter adjusted the test weight by a minimum of 100 g in the direction specified by the subject to make the test weight feel the same as the reference weight. This did not prevent the experimenter from overshooting the reference weight as the size of the adjustment varied. This procedure was repeated until the test weight was reported to be the same as the reference weight. Ten matches were made.

The weight matching protocol was repeated later with a 20 kg weight placed over the knee and distal thigh of the reference leg so that its gravitational vector passed through the knee and ankle joints. This arrangement increased the normal ground SENSATION OF STANDING 57

reaction force on the reference foot without affecting the net torque required to lift the reference weight.



Figure 3.3 Pressure under the feet A. Subjects were seated while they matched weights lifted by plantarflexion with increased normal thrust force created by a 20 kg weight placed over the distal thigh of the reference leg or without it (control). B. Individual data for ten comparisons and C. group mean data (± S.E.M.) normalised to the reference weight (dashed lines) show no changes in the mean or accuracy of matching.

Individual and group data (Figure 3.3B&C) show no effect of this normal force on the mean level matched or the accuracy of matching. The mean perceived heaviness of the lifted 2.0 kg reference weight was 2.0 kg [1.7, 2.3] with and without the 20 kg normal load (P = 0.97 by repeated-measures ANOVA).

Conclusion. Ground pressure on the feet does not affect the perception of actively applied force, making it unlikely that it masks the perception of force applied during standing. Study 3: Passive tension?

In the posture of normal standing, significant passive torque is exerted through the ligaments and tendons at the ankle. Study 1 showed that at the angle of standing for this group and situation, approximately half of the balancing torque arose through passive stiffness mechanisms. The effect of this passive torque on the perception of force exerted is investigated here. Subjects reported the perceived force exerted during SENSATION OF STANDING 58

standing by a voluntary contraction to replicate it while seated with the ankle in three different positions that markedly change the torque produced by passive ankle stiffness. Seven subjects were studied.

Subjects stood on the force platform and attended to the torque they exerted against the floor (as in Study 1). They then sat quickly in a rigid chair with the knees extended and feet on foot plates that were fixed in a pre-set position. This took 5 – 10 s. They then produced a voluntary plantarflexion contraction to match the force perceived during standing. In separate trials, subjects reported this force with the ankles at the angle that had been measured with a goniometer while standing, 5° more dorsiflexed than standing, and 5° more plantarflexed than standing (Figure 3.4), thus varying the passive ankle stiffness and the resting muscle length. Ankle torque was recorded from force transducers fastened to the foot plates. Ten matches were made.

The passive ankle stiffness was measured by strapping the body upright to the rigid support with the feet on coupled torque plates that had an axis of rotation through the axis of the ankle (Figure 3.4A). Torque applied to the plate was measured over 5 s epochs with the ankle stationary at different angles and the leg muscles at rest, which was monitored by soleus and tibialis anterior EMG bilaterally. To account for subjects’ different physiques, torque was normalised to the load stiffness of each subject with the zero angle aligned at the point of zero torque. Load stiffness was measured here as the slope of the torque-angle relationship while subjects swayed very slowly (i.e. quasi static) about the ankles through wide excursions (see Fitzpatrick et al. (1992) for method). In a single subject, ankle torque was recorded as the upright body was passively supported at different angles while the feet were fixed on a force plate (setup in Figure 3.1B).

Passive ankle torque increased with dorsiflexion (Figure 3.4B) and, in agreement with Study 1, was approximately half of the total torque at the typical lean angle of standing (~4 – 6°). The mean slope of the passive torque-angle relationship shows that a 10° excursion in ankle angle changed passive torque by ~80% of the total torque that subjects normally used to stand. Thus, this property is used to test the effect of passive stiffness on the perception of total torque applied during standing by having subjects SENSATION OF STANDING 59

report perceived force during standing with the ankle in different positions across this 10° range (Figure 3.4C&D).



Figure 3.4 Effects of passive tension A. Passive torque was measured by supporting the subjects upright and rotating the ankles to different angles () with the leg muscles relaxed. B. Passive torque increases markedly as the body leans forward and the ankle becomes more dorsiflexed. Mean passive stiffness is shown for 7 subjects with the angle measured relative to the point of zero net torque during standing (tdc: top dead centre). Torque is normalised for each subject relative to the measured load line that defines the equilibrium torque necessary to balance the body at any angle. Blue zone represent the region of normal standing. Note that at this point, active plantarflexion (ap) is necessary but at tdc a dorsiflexion (ad) contraction is require to overcome the passive ankle stiffness. C. Passive torque recorded in an individual subject supported upright at different ankle angles. Vertical arrows represent the normal angle of standing ± 5º, the points at which perceived force is reported in this study. D. Subjects initially stood and considered the force they exerted against the floor before being seated to reproduce that force. The ankle was pre-set at three different angles: normal angle of standing, five degrees dorsiflexed (DF) from standing or 5º plantarflexed (PF) from standing. E. Group mean (± S.E.M.) perceived active torque at different ankle angles as a percentage of the active torque.

Figure 3.4E shows that when seated with the ankle at the natural angle of standing, subjects underestimated standing ankle torque. Mean perceived active torque was 28.2% [20.8, 35.6] of exerted active torque and not different to the perceived ankle SENSATION OF STANDING 60

torque matched when upright (P = 0.75 by ANOVA). When the ankle was dorsiflexed and plantarflexed by 5º from the normal standing angle, active ankle torque was perceived to be 27.0% [20.7, 33.3] and 30.8% [24.0, 37.6] of exerted active torque respectively, differences not statistically significant by ANOVA.

All subjects were surprised when the experimenter moved the ankles to the angle of standing while they were sitting, reporting that it felt much more dorsiflexed and that they were very aware of the compression pressure exerted by the foot plate and tension in the calf muscles. They did not make this observation when the ankles were at the same angle when they were supported upright with the leg muscles relaxed.

Conclusion. Passive ankle stiffness generates a significant proportion of the torque that balances the body during standing. While this reduces the level of active torque required to stand, it does not interfere with the perception of the total applied torque. It appears however that the passive component during standing is not perceived to the same extent as during sitting. This suggests a masking of the passive torque, but not the active torque, by the large normal thrust force through the feet. Study 4: Self balance?

This experiment tests the hypothesis that the perception of force exerted is altered when balancing our own body compared with making an equivalent contraction to balance a non-self body. An inverted pendulum was used to simulate the human body during standing (Figure 3.5A; see Fitzpatrick et al. (1992) for detailed method). Seven subjects were studied.

A rigid post supported the subjects as they balanced an inverted pendulum that required the same contraction force previously recorded during normal standing. Subjects were permitted 10 s to consider the force of contraction before the pendulum was locked in place. Once the subjects were completely relaxed, as monitored by soleus EMG, they reproduced the perceived force of contraction in the calf muscles. Ten matches were made. Ankle torque was recorded from torque transducers that coupled the foot plates to the pendulum. SENSATION OF STANDING 61



Figure 3.5 Self balance? A. Subjects were secured firmly to a rigid support as they balanced an external pendulum with their feet. B. Raw data for a single trial. The subject was balancing the pendulum for 10 s before the foot plates and load were locked (first vertical line), and when the subject was relaxed, they matched the ankle torque used to balance the pendulum (second vertical line). C. Group mean data (± S.E.M.) of the perceived active torque as a percentage of the exerted active torque.

Data from one trial of an individual subject (Figure 3.5B) show that ankle torque was accurately perceived. The active torque required to balance the pendulum was 16.4 Nm and this was matched with an active perceived torque of 16.7 Nm. Although not statistically significant, mean perceived active torque was again (see Study 1, Figure 3.2) overestimated by 5.3% [-4.4, 15.0] compared to the active torque generated to balance the pendulum (Figure 3.5C).

It is worth noting here that subjects felt comfortable with the forces acting on the feet and did not report that it felt particularly heavy or that they had to make excessively forceful contractions. Although the ankles were at the angle of standing, identical to that of Study 3, they did not report the same excessive pressure from the footplates and tension in the calf muscles, even though this would have been significantly greater here with the added active contraction. This again indicates that the passive component of the applied torque, here 54% of the total torque, does not generate a strong perception. SENSATION OF STANDING 62

Conclusion. For this muscular contraction equivalent to that of standing in terms of torque and load stiffness, unlike standing, the force exerted is perceived accurately. As sensory masking phenomena have been excluded by the previous studies, this implicates the neural processes of muscle activation in the altered force perception during standing. Study 5: Cortical and sub-cortical contributions

The purpose of this experiment was to determine whether the underestimation of force during standing is related to contributions of cortical and sub-cortical drive to the leg muscles. Cortical involvement in a range of tasks is assessed by corticomuscular coherence and sub-cortical involvement is assessed by galvanic vestibular stimulation.

Corticomuscular coherence This study required measurable levels of corticomuscular coherence between electroencephalographic (EEG) and electromyographic (EMG) activities during a sustained muscle contraction. This cannot be obtained in all subjects for contraction of lower-limb muscles. Thus, subjects were first screened by having them maintain a 3-minute seated contraction of the calf muscles at the level previously recorded during standing. Of 15 subjects screened, the 10 who showed corticomuscular (EEG-EMG) coherence above the 95% confidence interval participated in the study.

Corticomuscular activity was recorded as subjects stood freely on a force platform (standing) and when they voluntarily contracted the calf muscles while supported upright by the rigid post (support), as illustrated in Figure 3.1A&B. Both began with the subject at rest and passively supported in the standing posture for 3 minutes. In the standing trial, the rigid post was removed to allow free standing for 3 minutes. In the voluntary trial, the subject remained supported by the rigid post as they contracted the calf muscles for 3 minutes to a displayed target torque set at the mean level measured during standing. The standing and support trials were performed in random order with rest periods between.

EEG was recorded from 14 surface electrodes placed at the vertex (CZ) and at 2 cm intervals from the vertex as arranged in Figure 3.6A. The common reference electrode was 6 cm frontal of CZ. EEG signals were amplified, bandpass filtered (0.3 Hz – 300 Hz) and recorded with 18 bit resolution at 2 kHz. EMG was recorded SENSATION OF STANDING 63

from surface electrodes placed 6 cm apart and oriented longitudinally over the soleus muscles bilaterally, amplified and bandpass filtered (30 Hz – 1 kHz).



Figure 3.6 Corticomuscular coherence for an individual subject A. Positions of the fourteen recording electrodes on the scalp relative to the vertex with an inter- electrode distance of 2 cm. Corticomuscular coherence is shown for a single subject during the standing (B) and supported conditions (C), in the layout of the electrode placements. Data for the right soleus is shown in blue and the left soleus in black, with pink lines indicating the 95% confidence limit. Coherence between the cortex and each soleus muscle were greatest at the vertex, shown in the box.

Using the method described by Rosenberg et al. (1989), coherence and cumulant density functions were constructed from the EEG and rectified EMG signals with EEG as the reference (input) signal. Across subjects, the EEG and EMG data for each subject were concatenated to create a single pooled data array for each condition. Coherence functions were calculated using a window of 4096 points (~2 s) giving a frequency resolution of 0.49 Hz. Frequency-specific coherence and time-specific cumulant density were deemed significant when values exceeded the 95% confidence limits described by Halliday et al. (1995).

Data for a single subject show that coherence between EEG and EMG activity was greatest at the vertex for the standing and supported conditions (Figure 3.6) with significant peaks at 12 – 30 Hz. Correlated EEG-EMG activity during standing was weaker than the supported condition for all corresponding scalp electrodes. SENSATION OF STANDING 64



Figure 3.7 Pooled corticomuscular coherence Corticomuscular coherence for the group is calculated from the concatenated data of all subjects, and cumulant density functions are their transform into the time domain. Both functions are shown for the standing (A) and supported conditions (B) for the Cz (vertex), C3 (left sagittal) and C4 (right) electrodes. Data for the right soleus is shown in blue and the left soleus in black, with pink lines indicating the 95% confidence limit.

Pooled coherence showed the same behaviour across subjects (Figure 3.7). Coupled EEG-EMG activity was greatest at the vertex in both conditions. Peak pooled coherence at the vertex during standing was 0.018 for the left soleus, at 25.2 Hz, and 0.015 for the right soleus, at 27.5 Hz. In the supported condition, peak coherence was 0.042 for both muscles, at 24.3 Hz. Pooled coherence estimates above the 95% confidence interval for standing represent 38.8% (left soleus) and 31.1% (right soleus) of the coherence for the supported condition (combined mean = 35.0%). Pooled coherence spectra at C3 and C4 show a clear asymmetry for the right and left soleus muscles, the stronger coupling being contralateral.

Cumulant density functions reflect the same pattern of association between EEG and EMG activities in the time domain for the standing and supported conditions. Significant peaks and troughs were observed at the same latencies in both conditions. SENSATION OF STANDING 65

For the vertex EEG-EMG pairs, significant peaks were separated by 42 ms, which translates to a peak coherent frequency of 23.8 Hz.

Vestibular responses Drive to the leg muscles from non-cortical centres might not contribute to the sense of exerted force during standing. Thus, this study determines the strength of associations between the vestibular (sub-cortical) drive to the plantar flexor muscles (i) during standing, (ii) during balancing the inverted pendulum and (iii) when performing a volitional contraction with the body supported upright.

Reflex responses to galvanic vestibular stimulation reflect the strength of the drive to active muscles that is under vestibular influence (Fitzpatrick & Day, 2004). The transmastoidal stimulus bypasses the sensory organs and introduces a pure vestibular disturbance by modulating the firing rates of primary vestibular afferents. In the leg muscles during standing the stimulus evokes short-latency (~55 ms) and medium- latency (~100 ms) responses (Cathers et al., 2005).

Ten subjects were studied with the 3 test conditions performed on separate days. In each, they had the eyes open with the head upright and turned to the right by 90° (Figure 3.8). In this posture, galvanic-evoked vestibular responses are maximised (Lund & Broberg, 1983; Cathers et al., 2005). The feet were side by side and ~20 cm apart. In the supported condition they maintained the contraction level of standing by viewing a target on an oscilloscope, which was extinguished for the periods of the galvanic stimuli.

Bipolar currents from a controlled current source were delivered through 3 cm2 Ag-AgCl electrodes adhered over the mastoid processes. Rectangular current waveforms (3 mA 2 s) were delivered with randomised between-stimulus intervals of 3 – 5 s. The on and off transients were sigmoidal over 3 ms to reduce cutaneous sensation. Stimuli were randomised for polarity (anode left, anode right) and presented in 2 blocks of 32 with rests between. EMG was recorded bilaterally over soleus and bandpass filtered (30 – 1 kHz, 30 – 300 Hz for the standing condition). EMG records, 1 s before and after stimulus onset, were normalised for each subject to the mean RMS amplitude of the group before rectifying and averaging for each stimulus polarity across trials and subjects. The anode-left responses were inverted about the pre-stimulus mean SENSATION OF STANDING 66

level and combined with the anode-right responses to create final averages of 64 bilateral responses.

 Figure 3.8 Galvanic-evoked vestibular responses Vestibular reflex responses in soleus were sought by rectangular wave, bipolar galvanic stimulation while standing (A), while supported upright and making an equivalent contraction of the calf muscles (B), and when balancing the inverted pendulum matched to the body (C). Strong biphasic responses observed when standing were (almost) non-existent in the supported and pendulum conditions.

The difference in responses between these conditions was dramatic, with very strong biphasic modulation of ongoing soleus EMG activity during standing but almost no responses in the supported volitional contraction and when balancing the inverted pendulum, despite baseline EMG levels having been matched closely. SENSATION OF STANDING 67

Conclusions. Corticomuscular coherence indicates that the strength of cortical association with leg-muscle activation during standing is approximately one-third that when making a volitional contraction with the same force output. The presence of vestibular reflex activity only during standing indicates that this drive to the leg muscles is only activated during standing. The corticomuscular coherence follows the pattern of force perception in Study 1, with both coherence and perception reduced to approximately one-third during standing. In contrast, the vestibular responses are the inverse of the level of active force exerted in Studies 1, 3 and 5. This suggests that much of the active force exerted during standing is not perceived because its motor drive arises from a sub-cortical system to which the cortex does not gain perceptual access. General discussion

This study investigates the common observation that standing is not perceived as particularly difficult or requiring a forceful contraction. The torque applied to balance the body during standing was perceived to be approximately one-third of an equivalent voluntary contraction that did not involve balancing the subject’s own body. The results show two reasons for this reduced perception of force during standing. First, passive mechanisms account for a significant proportion of the standing ankle torque and, second, the active muscle contraction originates predominantly from sub-cortical control processes.

I will consider these observations in terms of the reafference principle concept of von Holst and Mittelstaedt (1950, 1973) in which sensory afferent signals contain a reafference signal that represents the consequences of motor actions, and an exafference signal that is the consequence of external influences. The neural processes that separate the two involve the motor command or its efference copy. Chapter 2 shows that this efference-copy process begins at the level of the muscle and muscle spindle for the perception of exerted force.

Passive mechanisms. Approximately half of the ankle torque that supports standing is derived from passive mechanisms. Unlike most muscles the normal operating range for the plantar flexors does not correspond to their resting muscle lengths (Sale et al., 1982). In fact, the optimal range for the ankle joint, in terms of force production, corresponds with the more dorsiflexed posture adopted in standing, suggesting an SENSATION OF STANDING 68

evolutionary adaptation of the passive mechanisms in addition to the muscular for our normal upright posture.

Study 3 reveals that passive tension at the dorsiflexed angle of standing does not interfere with the perception of the actively applied torque during standing. This is consistent with the conclusion of Chapter 2 that force perception is largely based on a reafference signal, if we accept that passive mechanisms do not generate a reafference that can be extracted from the total afferent signal.

Passive torque can be perceived as all subjects reported that when the ankles were passively positioned at the angle of standing while they were sitting, they were aware of the pressure exerted by the foot plate, the tension in the calf muscles, and that the ankles felt more dorsiflexed than when standing. In addition to showing that joint position sense is not absolute but intimately associated with local passive forces, this observation indicates that torques exerted by passive and active processes during contraction are handled differently by perceptual processes. This may explain why subjects were not aware of the passive torque in the three upright conditions when the calf muscles were active (i.e. standing, supported upright, and balancing the pendulum) even though the total pressures and forces were higher because of the contraction. However, the absence of this sense of passive force is unlikely to be associated with the contraction as it was also absent when they were supported upright with the muscles at rest. Thus, it appears that the large forces through the feet and ankles from the weight of the body mask the sense of passive force. Overall, the simplest explanation is that passive torque is extracted as exafference and interpreted as something the world is doing while active torque is extracted as reafference and interpreted as something the actor is doing.

Why then is the contraction made during standing only perceived as one-third that of a volitional contraction of the same force?

Control of standing. Evidence that motor activity in the calf muscles during standing is predominantly driven by sub-cortical systems comes from the results of Study 5. The implication is that there is no conscious perception of the effort or force generated by the sub-cortical drive and we can speculate that there are distinct cortical and sub- cortical processes concerned with balance control. SENSATION OF STANDING 69

Corticomuscular coherence was strongest at the vertex, the sensorimotor area of the cortex that is associated with the lower limb. During standing, coherence was reduced to approximately one-third that when making a volitional contraction with the same force output. Corticomuscular coherence represents the strength of association between oscillatory activity in the cortex and the muscle and is interpreted to reflect changes in corticospinal drive to the muscle (Perez et al., 2006). The 12 – 30 Hz bandwidth of synchronous cortical and muscle activity measured in this study is within the frequency range that corresponds with maintained cortical activation of muscles (Conway et al., 1995; Salenius et al., 1997; Kilner et al., 1999). This suggests that the cortical motor command to the calf muscles during standing is relatively small.

An alternative view proposes that corticomuscular coherence reflects sensorimotor integration (MacKay, 1997). That is, coherent activity between the cortex and the muscles depends not only on corticospinal drive but also reafferent feedback that completes a sensorimotor loop. This view is supported by the indirect evidence that activity in peripheral afferents recorded from dorsal root ganglia in awake monkeys is correlated with muscle activity over the same frequency range identified for corticomuscular coherence (Baker et al., 2006). The psychophysical experiments in Study 1 support this concept of corticomuscular coherence where both coherence and perceived force, which I propose is based on reafference, were reduced to approximately one-third of the force-matched voluntary contraction.

Evidence that the dominant sub-cortical drive to the calf muscles during standing receives a major vestibular input comes from the galvanic-evoked muscles responses. These reflexes are specific to standing balance and are not present when the same force output of standing is generated by volitional effort in situations other than free standing.

The vestibular reflex picture is the complement of the corticomuscular coherence picture. Standing recruits a strong vestibular sub-cortical descending drive and leaves a smaller (approximately one-third) cortical drive whereas volitional contraction only recruits cortical drive. The reafference that provides for cortical perception is extracted from the total afference by reference to the cortical efference copy signal. Thus, the explanation for the weak perception of force exerted during standing is that the cortex identifies its own actions by extracting reafference but cannot extract the sub-cortical SENSATION OF STANDING 70

reafference because it does not possess the sub-cortical efference copy. In this regard, it is worth noting that postural responses evoked by galvanic stimuli are irresistible and not overridden by volitional intent even when the subject delivers them repeatedly (Guerraz & Day, 2001, 2005), and subjects report these movements in their actual directions unaffected by the evoked vestibular signal of movement in the opposite direction (Fitzpatrick et al., 1994a). This behaviour is as if the cortex observes the outcome of an independent sub-cortical motor drive and in effect treats it as exafference.

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Chapter 4

Perfusion and lower-limb muscle function

Chapter 3 shows that much of the neural drive for standing is sub-cortical and bypasses the efference-reafference mechanism of force perception identified in Chapter 2. Like this perceptual process, blood pressure regulation during muscular work also relies on a motor efference copy driving cardiovascular centres to generate a “central pressor response”. This leads me to identify the action of this pressor response during standing as the sub-cortical motor drive might likewise bypass the central pressor response. However, the muscle metaboreflex might also evoke a pressor response and the leg muscles are uniquely exposed to high perfusion pressures during standing. Identifying the central pressor response requires an understanding of the influences of perfusion pressure on the contractility of the leg muscles and the influence of the metaboreflex from these muscles. Thus, this study investigates how local perfusion pressure and central arterial pressure affect tibialis anterior force output during electrically-evoked (no central pressor response) and voluntary contractions. Changing the height of tibialis anterior relative to the heart controlled local perfusion pressure. Stimulated force output was highly sensitive to variations in perfusion pressure across the physiological range. This perfusion-dependent change in muscle contractility occurs quickly (53 s time constant) and can be explained by the ionic shifts that mediated changes in membrane potential. These effects are reversible and indicate a steady-state equilibrium between force output and perfusion pressure. These stimulated contractions, which approximated the workload of standing, did not produce significant cardiovascular responses, indicating that the muscle pressor reflex does not play a major role in cardiovascular regulation at this workload. Voluntary contractions with force output that would require a constant motor drive at steady perfusion generated a central pressor response when perfusion pressure was lowered as a larger drive compensated for the lost contractility. The relationship between muscle contractility and perfusion for this large postural muscle was consistent with that of a small hand muscle, and produced similar effects for passive or active changes in perfusion pressure. MUSCLE PERFORMANCE 72

Introduction

Co-ordination between motor and cardiovascular control is needed to maintain muscle performance. Regulatory processes that decrease local vascular resistance and raise arterial blood pressure attempt to match muscle blood flow to the metabolic demands of contraction (Andersen & Saltin, 1985).

Relatively small changes in arterial pressure across the physiological range of blood pressures (75 - 125 mmHg, Master et al., 1951) result in proportional changes in muscle blood flow and, as a consequence, tetanic force output from the stimulated muscle (Hobbs & McCloskey, 1987). Thus, local autoregulatory mechanisms that enhance muscular blood flow by modulating the resistance of the vascular bed during contraction do not compensate for changes in perfusion pressure. In a human hand muscle, decreasing muscle perfusion pressure within this physiological range reduces contractile force output (Fitzpatrick et al., 1996b) and as a result increases the rate at which the muscle fatigues with sustained voluntary contractions (Wright et al., 1999). Increasing systemic blood pressure partially offsets this decline in muscle performance (Wright et al., 2000).

Two main processes act to increase blood pressure during muscle contraction. As the motor command is sent to the muscles, a centrally-generated pressor command is sent to cardiovascular centres to raise blood pressure through an anticipatory or feedforward control process. From the working muscles, the metabolic and mechanical consequences of contraction signalled by muscle afferents also act on medullary cardiovascular centres to increase blood pressure in a reactive or feedback control process described as the exercise pressor reflex (Mitchell et al., 1983). This response is mediated by a muscle metaboreflex with an afferent arc stimulated by muscle chemoreceptors, and a muscle mechanoreflex stimulated by mechanoreceptors sensitive to muscle stretch (Paintal, 1960; Kaufman et al., 1983; Mense & Stahnke, 1983).

The central pressor response is approximately proportional to the intensity of the motor command rather than the force output of the driven muscle (Lind & McNicol, 1967b). In other words, commands that maximally activate large and small muscles will produce similar increases in blood pressure. For the intrinsic hand muscles, voluntary contraction results in robust increases in systemic blood pressure whereas an equivalent MUSCLE PERFORMANCE 73

workload generated by electrically stimulated contractions has no significant effect (Fitzpatrick et al., 1996b; Wright et al., 1999). This means that for these small muscles the central pressor response is responsible for practically all of the increase in central blood pressure while the peripheral metaboreflex does not contribute significantly.

Unlike the central pressor response, the metaboreflex is approximately scaled by the amount or volume of muscle activated (McCloskey & Streatfeild, 1975). As it is driven by the metabolic consequences of contraction, the balance between work output and muscle perfusion determines the rise in blood pressure. The unique upright human posture means that our limb muscles are exposed to a wide range of perfusion pressures through the effects of addition or subtraction of the hydrostatic pressure head that occurs during limb elevation and body posture (Nielsen, 1983; Fitzpatrick et al., 1996b; Gamble et al., 1997). These effects are likely to be critical for the normal function of the human leg muscles during postural control.

Apart from being large muscles, the leg muscles normally work continuously rather than intermittently and for prolonged periods in maintaining upright posture and in this role they are exposed to a large and constant hydrostatic pressure. Our ability to maintain an upright posture for extended periods suggests special adaptations of these muscles. Vascular reactions to a stimulus vary between the arm and leg, with the arm showing greater levels of vasodilation (Newcomer et al., 2004) and the leg more sensitive to sympathetically-mediated vasoconstriction (Pawelczyk & Levine, 2002). Through their size, intrinsic vasculature, high perfusion pressures and continuous action in postural activities, the leg muscles are likely to respond differently to changes in perfusion pressure and be associated with different pressor responses, making it difficult to extrapolate the behaviour of the small hand muscle.

This study investigates the relationships between muscle force, perfusion pressure, and the pressor responses evoked by the contracting tibialis anterior muscle (TA), a predominantly oxidative type muscle (Henriksson-Larsen et al., 1983) used in balance control (Fitzpatrick et al., 1994a) with a fibre-type composition similar to the soleus muscle (Johnson et al., 1973). Local perfusion pressure is controlled by passively raising or lowering the legs during electrically-evoked and voluntary contractions. The effect of central blood pressure on muscle force output is investigated by increasing MUSCLE PERFORMANCE 74

arterial blood pressure by contraction of the forearm muscles. Cardiovascular responses during evoked and voluntary contractions are compared to determine the influences of the central pressor response and the exercise pressor reflex on blood pressure. Methods

Experiments were conducted on 8 healthy subjects (5 males), aged between 26 and 62 years (mean 35.8).

Setup Subjects were studied in two different positions using the same apparatus (Figure 4.1). Each heel rested on a stable support that allowed both legs to be passively raised or lowered by rotation about a horizontal axis aligned through the hip joints. This arrangement ensured that the ankle remained at the same angle (90°) as the leg was raised or lowered. The right foot was strapped securely to a foot plate that had an axis of rotation aligned with the ankle but immobilised by an isometric load cell from which dorsiflexion ankle torque was measured.



Figure 4.1 Method The tibialis anterior muscle was positioned at heart level at the start of each trial. In A, both legs were passively raised (+H) by the experimenter until the muscle was 30 cm above heart level. In B, both legs were passively lowered (-H) until the muscle was 30 cm below heart level. Tetanic contractions were produced by supramaximal stimulation over the motor point. Muscle force output (torque), central and leg arterial pressure were measured continuously during each trial.

Tetanic contractions of the right tibialis anterior muscle were evoked by monopolar stimulation over the motor point of the muscle using a voltage-controlled isolated current source. Surface electrodes (5 3 cm) were placed approximately 3 cm below the tibial tuberosity and separated by 4 cm. Tetanic contractions were produced, MUSCLE PERFORMANCE 75

once every second, by a train of five supramaximal pulses of 1 ms duration at 25 Hz. Supramaximal stimulation was established by identifying the stimulus strength that produced a maximal single twitch force and then increasing the stimulus by 20% to ensure maximal activation. This stimulation protocol was chosen to ensure that the contractile responses in fatigue-resistant muscle fibres would be preserved for the duration of the stimulation protocol (Burke et al., 1973) and because the workload is not dissimilar to the 10 – 20% of maximal activation of the calf muscles during standing (Okada, 1972).

Arterial pressure and heart rate were measured continuously by digital plethysmography (Portapres Model-2: Finapres Medical Systems, Amsterdam). Central arterial pressure was measured in the resting middle finger, correcting for the hydrostatic pressure of a column of blood between the heart and the finger. The muscle contractions prevented direct measurement from the stimulated leg so local arterial pressure was measured in the second toe of the left, non-stimulated leg, which was raised and lowered in parallel with the right leg. Arterial pressure of the stimulated tibialis anterior was estimated by adding the hydrostatic pressure between muscle centre and the toe.

Protocol Subjects were instructed to refrain from consuming alcohol, caffeine and strenuous exercise for at least 4 hours before the experiment. Before each trial, subjects rested for 10 minutes in the starting position to ensure stable baseline conditions. For experiments in which local perfusion pressure was reduced, subjects began supine with both legs extended and resting at heart level (Figure 4.1A). For experiments in which local perfusion pressure was increased, subjects began in a reclined, seated position with the legs extended and the hip flexed so that the tibialis anterior muscle was at heart level (Figure 4.1B).

Study 1: Effects of posture on arterial pressure in the legs. Arterial pressure in the toe was measured at different leg heights. With the subject at rest, both legs were raised or lowered to different positions that placed the feet between 120 cm below the heart level to 40 cm above heart level (measured as the height of the suprasternal notch). The feet were held at each position for 2 minutes. MUSCLE PERFORMANCE 76

Study 2: Effects of local perfusion pressure on muscle force output. The legs-raised (Figure 4.1A) and legs-lowered (Figure 4.1B) protocols were performed on different days in randomised order allowing at least two days between tests. During each trial, tibialis anterior was stimulated to produce a tetanic contraction every second for 600 s. Trials began with the muscle at heart level. After 240 s, the legs were raised or lowered by 30 cm to decrease or increase perfusion pressure, respectively. After 180 s at the test height, the legs were returned to heart level for the remaining 180 s of stimulation. Subjects were instructed and encouraged to remain relaxed throughout stimulation and particularly during raising and lowering the legs, which typically lasted ~10 s.

Study 3: Effects of the motor command on blood pressure. To determine the role of the motor command in restoring local changes in leg arterial pressure, the legs-raised and legs-lowered protocols described above were repeated but with voluntary rather than stimulated contractions. During the first 240 s, when the muscle was at heart level, subjects contracted the tibialis anterior at a workload set to be equivalent to the stimulated contractions, determined from the force-time integrals. The target force, presented to the subject on an oscilloscope, was matched to the steady decline in tetanic force recorded during the stimulated contractions. Voluntary contractions were produced every 10 s, contracting 3 s and resting 7 s, a duty cycle equivalent to that measured during the stimulated contractions.

After 240 s, the legs were raised or lowered. At this time, the target force was set to be the exponential extrapolation of the decline in force output during the first 4 minutes rather than tracking the increase or decrease in tetanic force that occurred during the stimulated contractions. Setting the target force on this trajectory maintains (approximately) a constant motor command for the muscle at heart level as this is the force output that would have been produced by constant tetanic stimulation. In other words, setting a constant force target would require an increasing motor command as muscle performance declined3. Thus, this protocol reveals the pressor responses that

 3 Note that this protocol differs from most previous studies [e.g. Eiken (1987), Egana & Green (2005)] that had subjects exercise at a constant or incremental workloads, a protocol that would see any potential loss of muscle force partially reconciled by an increased voluntary drive and the associated increase in central blood pressure produced through the central pressor response (Wright et al., 2000). MUSCLE PERFORMANCE 77

accompany the change in motor command required to offset the change in muscle contractility as perfusion pressure is reduced (or increased) by elevating (or lowering) the muscle.

The force generated by a maximum voluntary contraction (MVC) of tibialis anterior was measured on a separate day in the same apparatus with the leg at heart level.

Study 4: Effects of central arterial pressure on muscle force output. The effect of increased central arterial pressure on muscle force production was determined in 6 subjects by using a voluntary static handgrip contraction to increase central blood pressure. Using the protocol described above, tibialis anterior was electrically stimulated for 600 s with the subject supine and the muscle at heart level. After 240 s of stimulation, subjects used a hand dynamometer to make a steady handgrip contraction at 20% of their previously measured maximal force for 180 s, and then relaxed their hand for the remaining 180 s of evoked contractions.

Measurement and analysis Data were sampled at 50 Hz and recorded to computer. For each tetanic contraction, net force was measured as the difference between the pre-stimulus force on the foot plate and the maximum tetanic force achieved. To compare effects across subjects and conditions, forces averaged over 30 s epochs were normalised to the net force generated by the last contraction before each test condition was applied (i.e. before raising or lowering the legs, or starting the handgrip contraction).

Between- and within-subject differences in arterial pressure were corrected by normalising mean arterial pressure (diastolic + 1/3 pulse pressure) to the average arterial pressure across subjects during the last contraction before each test condition was applied. Mean arterial pressure was averaged over the same epochs as the force output. A similar correction was applied to measurements of heart rate before averaging over the same epochs as the force output.

Repeated-measures ANOVA was used to examine the significance of differences in mean arterial pressure (MAP) and heart rate (HR) during postural adjustments alone and for each test condition. Post-hoc Bonferroni corrections were made for multiple MUSCLE PERFORMANCE 78

time comparisons, with the level of significance set at P = 0.05. Group data are presented as means with 95% confidence intervals. Results Posture and arterial pressure in the legs With the legs at heart level, resting central mean arterial pressure was 74.9 mmHg [70.9, 78.9] and leg mean arterial pressure was 86.8 mmHg [79.7, 93.9] (Figure 4.2). At +30 cm, it was 27.1 mmHg [20.5, 33.7] lower and at -30 cm it was 30.8 mmHg [24.2, 37.3] higher. This relationship was highly linear and essentially as predicted by the difference in hydrostatic pressure, with a slope of -0.90 mmHg.cm-1. Central blood pressure increased slightly (3.0 mmHg.m-1; R2 = 0.67 [0.57, 0.77]) as the legs were raised and there was a tendency for heart rate to decrease (2.5 bpm.m-1; R2 = 0.62 [0.51, 0.73]).

 Figure 4.2 Resting arterial pressure with the legs at different heights Mean arterial pressure (MAP) in the leg (red diamonds) and central arterial pressure (open) with the legs passively positioned at different heights relative to the heart. The MAP predicted by the hydrostatic pressure of a column of blood between the leg and the heart is represented by the broken black line. Heart rate is shown in blue. The open circle (S) represents the arterial pressure in tibialis anterior during standing. Data are group means ± 2 S.E.M.

Local perfusion pressure and muscle force output Stimulated contractions of tibialis anterior produced peak tensions that ranged from 2.7 – 6.1 Nm across subjects. These were equivalent to 10 – 17% of individual maximum contraction levels. After an initial potentiation (20 s), force output followed a negative exponential decline with a prolonged time constant of 240 minutes (S.E.M. 18). Data for a typical subject are shown in Figure 4.3. Elevating the contracting tibialis MUSCLE PERFORMANCE 79

anterior above heart level resulted in an immediate reduction in muscle force. The opposite effect was seen when the muscle was lowered below heart level, force output immediately increased above the steadily declining baseline.

 Figure 4.3 Force of tetanic contractions of tibialis anterior Raw records from a single subject is shown for three separate trials when the tibialis anterior was positioned at heart level (A), raised (B) or lowered (C) by 30 cm relative to the heart. Broken lines are the times when leg position was passively raised or lowered. Each peak represents one tetanic contraction produced by stimulation every second.

During the first 240 s of stimulation, central MAP tended to increase by 4.5 mmHg [0.0, 9.0] above resting levels to reach new stable baselines during tetanic contractions. Heart rate increased by 4.2 bpm [1.1, 7.3].

When perfusion pressure was reduced by elevating both legs, muscle force fell gradually and reached a new steady state of decline after 120 s (Figure 4.4). Raising the legs by 30 cm reduced leg MAP by 28.5 mmHg [20.5, 36.5]. This corresponded with a significant reduction in tetanic force, 20.5% [9.2, 31.9] after 120 s, compared with the predicted from the initial rate of decline (broken line, Figure 4.4). Conversely, increased perfusion pressure augmented muscle force. Lowering the legs by 30 cm increased leg MAP by 26.3 mmHg [20.9, 31.7] with a maximum increase in tetanic force of 14.9% [8.0, 21.8] at 120 s. After the peak increase, tetanic force declined steadily but remained elevated above predicted values. Central MAP and HR remained relatively stable MUSCLE PERFORMANCE 80

throughout the stimulation and no significant changes were observed when the legs were raised or lowered. Returning the legs to heart level after legs-raised or legs- lowered periods restored leg MAP. Force recovered over a similar time course towards the values predicted from the initial rate of decline.

 Figure 4.4 Tetanic force and mean arterial pressure Raising the legs above the heart (filled symbols, black line) decreased leg mean arterial pressure (MAP) and mean force. Conversely, lowering the legs below the heart (open symbols, blue line) increased leg MAP and mean force. Both legs were passively held in the raised or lowered position for 180 s (shaded area). The broken line represents the predicted decline in mean force extrapolated from the first 240 s of evoked contractions (heart level). Open circles represent the mean force output of the legs raised and legs lowered conditions. HR, heart rate.

Motor command and blood pressure Voluntary activation of tibialis anterior at this work level had no significant effects on baseline blood pressure during the first 240 s of voluntary contractions (central MAP increased by 1.7 mmHg [-2.6, 6.0]; leg MAP increased by 8.2 mmHg [-0.8, 17.2]). Similarly, there was no change in mean heart rate.

Data for one subject is shown in Figure 4.5. The force profile was the same for the legs-raised and legs-lowered condition and, for this subject, decayed exponentially with a long time constant of 144 minutes. Raising the legs above heart level decreased leg MUSCLE PERFORMANCE 81

MAP but increased central MAP. Lowering the legs below heart level increased leg MAP, however, no change was observed in central MAP.



Figure 4.5 Responses of one subject during voluntary contractions of tibialis anterior Contraction force is shown for the legs lowered protocol only; however, the same force profile was used in the legs raised condition. Passively raising the legs (filled symbols, black line) above the heart reduced leg MAP and increased central MAP. Lowering the legs (open symbols, blue line) below the heart increased leg MAP without changes in central MAP. Both legs were held in each test position for 3 minutes (shaded area). HR, heart rate.

Across subjects, raising the legs above the heart during the voluntary contractions decreased leg MAP by 18.1 mmHg [9.8, 26.4] and lowering them below heart level increased it by 36.9 mmHg [20.1, 53.2] (Figure 4.6). A small increase in central MAP was observed after the legs were elevated (5.2 mmHg [2.4, 8.0] at 120 s) but there was no change when the legs were lowered. Heart rate remained steady throughout the voluntary contractions for both leg positions. Returning the legs to heart level, after the legs raised or legs lowered conditions, restored leg MAP towards the baseline established during the first 240 s of contractions.

MUSCLE PERFORMANCE 82



Figure 4.6 Group data during voluntary contractions of tibialis anterior Passively raising the legs above the heart reduced leg MAP (black). Central MAP increased when the legs were raised and was significantly so after 2 minutes. Lowering the legs (blue) below the heart increased leg MAP without significant changes in central MAP. Both legs were held in each test position for 180 s (shaded area). P < 0.05 when compared to values at T = 0. MAP, mean arterial pressure; HR, heart rate

Effects of raising central arterial pressure Central arterial pressure and heart rate increased steadily during the handgrip contraction at 20% MVC. After 180 s, HR increased by 8.7 bpm [5.4, 12.0] and central MAP and leg MAP increased by 11.3 mmHg [6.3, 16.3] and 16.4 mmHg [10.7, 22.1], respectively (Figure 4.7).

During the time of the handgrip contraction, this increase in blood pressure corresponded with a 4.8% [0.8, 8.8] increase in the stimulated force output of tibialis anterior above the predicted exponential decline in force output. After ending the handgrip contraction, blood pressure and heart rate dropped quickly towards the baselines established during the first 240 s. Meanwhile, stimulated TA force output began to decline but remained elevated above the original predicted trajectory of declining force output.

MUSCLE PERFORMANCE 83

 Figure 4.7 Effect of central blood pressure on muscle force output A voluntary static handgrip contraction at 20% MVC (shaded area) was used to increase central blood pressure. Heart rate (HR), central MAP () and leg MAP () increased steadily during the handgrip contraction. The force output for evoked contractions of tibialis anterior is shown here on an expanded scale with the predicted decline in mean force output, extrapolated exponentially from the first 240 s of evoked contractions, represented by the red broken line.

Discussion

This study demonstrates a potent effect of local perfusion pressure on the force output of the working tibialis anterior muscle. A 55 mmHg change in leg arterial pressure produced a 36% change in muscle force output. Thus, any local autoregulatory mechanisms that act to buffer muscle blood flow against changes in perfusion pressure are inadequate to maintain force output, even across these changes within the physiological range. An effect of similar gain has been described for a small intrinsic muscle of the hand (Fitzpatrick et al., 1996b). However, with a hydrostatic pressure head of more than 80 mmHg during standing and locomotion (Figure 4.2), this perfusion-dependent change in muscle contractile performance holds greater significance in the lower limb. Extrapolating the changes in force output produced by the changes in hydrostatic pressure delivered here (Figure 4.4) indicates that contractile performance and fatigue resistance of the leg muscles relies strongly on their orthostatic perfusion pressure. In effect, we can stand only because we are standing. MUSCLE PERFORMANCE 84

For the purposes of this thesis, the experiments here investigate the exercise and central pressor responses in human postural muscles so that the operation of these cardiovascular reflexes during standing can be determined. The work is also important for understanding the processes limiting force output of working muscles, and I will first consider this issue.

Fatigue, contractility and their perfusion dependence Continual contraction of a muscle reduces the maximum force it can generate, a phenomenon encompassed by the general term: “fatigue”. Many processes can contribute to fatigue including an inability to generate a maximum command within the nervous system, loss of support or inhibition of the motoneuron from muscle afferents, neuromuscular junction failure, ionic changes in the muscle, metabolite delivery and removal, and processes at the level of the contractile mechanism (see Enoka & Stuart, 1992; Fitts, 1994; Gandevia, 2001; Allen et al., 2008 for reviews). Effects from the neuromuscular junction onward should be considered to explain the loss of force output in the study here that delivers supramaximal stimuli at the muscle motor point.

Repeated tetanic stimulation with constant perfusion pressure produced a long- term decline in muscle force output. At the relatively low workload of this study, the time constant of this decline was estimated at ~240 minutes, the same as that measured for adductor pollicis when stimulated at the same stimulation protocol (Fitzpatrick et al., 1996b). In contrast, the effect of perfusion pressure on force output operates with a short time constant, estimated here by graphical solution as ~48 s (Figure 4.4). Furthermore, it acts reversibly about the baseline established by the long-term decline in tetanic force. In the study by Fitzpatrick et al. (1996b), the decline in stimulated force output was best fitted by a double exponential decline with time constants 53 s and 240 minutes (shown in Figure 4.8A)4.

 4 The rapid decline at the start of the tetanic stimuli, which corresponds to the 53 s response, is not seen clearly in the tibialis anterior recordings here. The explanation we favour is that it was more difficult to stabilise the leg than the thumb and this resulted in a “settling in” period, which might include volitional adjustment as the stimuli were significantly more painful at the onset. It is also possible that it is a true muscle effect that could arise through viscoelastic change or early changes in the excitation-contraction processes. MUSCLE PERFORMANCE 85

The ~48 s time constant of this reversible change in force output and the 53 s reported by Fitzpatrick et al. (1996b) are equivalent to the 54 s time constant reported for the recovery of the sarcolemmal resting membrane potential (EM) in the mouse soleus after its depolarisation by repetitive stimulation (Juel, 1986). This was associated + with a doubling of extracellular potassium [K ] at the relatively intense stimulation + workload and a recovery of intracellular [K ] that paralleled the EM recovery. The shift + of EM in the direction of the sodium (Na ) equilibrium potential, brought about + predominantly by an increase in interstitial [K ] during muscle work, is thought to be a major cause of the contraction-related decline in force output through the effect it has + on reducing inward Na current, membrane excitability and conduction speed (Sejersted & Sjogaard, 2000; Cairns & Lindinger, 2008). In human gastrocnemius muscles, Green + et al. (2000) show that interstitial [K ] increases linearly with plantarflexion power output across the physiological range. Their results, obtained by microdialysis, show + that even at a relatively low power output of 13% maximum (0.9 W), interstitial [K ] rose from 4.5 to 5.7 mmol.l-1 but reached mean levels of ~12 at higher workloads, before recovery with a mean half-time of 57 s on stopping the exercise. Furthermore, these changes were associated with parallel increases in interstitial-arterial and + interstitial-venous K concentration gradients despite large increases in leg perfusion during the contraction period.

These considerations indicate that the reversible modulation of force output by perfusion pressure observed here is likely to be the result of these ionic shifts. In the + + resting muscle, interstitial [K ] approaches the theoretical lower limit of plasma [K ] + and in the muscle working at extreme workloads it moves towards intracellular [K ], + which declines during activation. The two K fluxes largely determine the interstitial + + [K ] equilibrium. Leak of K from myocytes through voltage-gated channels is + proportional to workload, and vascular removal of K is proportional to perfusion (Figure 4.8B). This simple model of interstitial ionic equilibrium and its effects on contractility can explain the reversible effects of altering muscle perfusion pressure observed here.

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Figure 4.8 Model of perfusion effects on force output A. Recording from Fitzpatrick et al. (1996) of force output from adductor pollicis at heart level stimulated with the protocol used here. Force output was fitted by a double exponential with time constants 53 s and 240 minutes, equivalent to the two time constants identified for tibialis anterior in the present study. B. A simple model of interstitial [K+] equilibrium determined by flux into the space from voltage-gated channels (vg) proportional to the workload, and flux out determined by energy-dependent, rate-limited uptake into the intracellular space, and by vascular removal proportional to perfusion (p) and interstitial [K+]. C. The scaling factors Q and P in the double exponential decline function represent the relative strengths of the short-term interstitial ionic equilibrium process (53 s) and the long-term fatigue process (14400 s) respectively. The thick orange curve (low workload) is the double exponential function fitted to the data of A. The thick blue curve represents a higher workload with Q and P scaled higher by an arbitrary factor. For both, the force curves shift up or down when Q is changed to represent different perfusion levels (high and low Q). At points x and y, Q is changed up and down and the broken lines show the transitions between the high-Q and low-Q trajectories. D. Model equation is fitted to data from tibialis anterior where the leg is raised then lowered.

A model of perfusion effects on force output is shown in Figure 4.8C for a muscle contracting at low and high workloads. The double exponential decay curve at a constant perfusion pressure reflects the separate short and long time-constant mechanisms at play. The short time constant represents the perfusion-contraction interstitial ionic equilibrium and the long-term decline a slower energy-dependent change in the excitation-contraction chain. At a low workload and constant perfusion pressure (orange curve), the short-term equilibrium is attained and fatigue proceeds on the long-term trajectory for that workload and perfusion pressure. Changing perfusion pressure (x) establishes a new interstitial ionic equilibrium point and force shifts MUSCLE PERFORMANCE 87

accordingly with the short time constant to a higher or lower long-term trajectory. At a higher workload (blue curves), the declines are amplified but the processes take the same time constants. Figure 4.8D is a best fit of the model double-exponential equation for contraction data from tibialis anterior. At the point of raising the legs, the gain constant (Q) of the perfusion term of the equation is increased, and on lowering the legs it is restored to the initial value. The lower curve shows the extrapolation of the high-Q function.

Effects of blood pressure on muscle contractility The reversible effect of central blood pressure on the force output of the working tibialis anterior was calculated to have a proportional gain of 2.1 in restoring force output to its initial pre-work maximum. That is, a 10% increase in central arterial pressure restored lost force output by just over 20%. A similar gain has been reported for the small hand muscle adductor pollicis (Wright et al., 2000). The comparison of these hand and leg muscle responses (Figure 4.9) suggests that these perfusion dependent changes in muscle contractility are consistent across a wide range of perfusion pressures, muscle groups and muscles sizes. It also indicates that passive (posture related) and active (contraction related) changes in perfusion pressure produce similar proportional effects on muscle contractility.

 Figure 4.9 Relationship between force output and blood pressure in working muscle The percentage change in stimulated force output calculated relative to the initial maximum is plotted against the percentage change in blood pressure for the different manipulations tested. The effects of passive (postural) changes in local perfusion pressure are shown for raising and lowering the legs as red and blue triangles, respectively. The effects of increases in blood pressure produced by active contraction of the upper limb are shown as the open diamonds. Regression of these data shows a gain of 2.13. The data of Wright et al. (2000) showing the response of adductor pollicis, plotted in green, show the same proportional response. MUSCLE PERFORMANCE 88

Other influences on contractility. Working muscles at moderate intensity (>25% MVC) increases systemic sympathetic nervous activity (Mark et al., 1985; Seals, 1989b). Any increase in leg vascular resistance through this sympathetic activity should decrease muscle blood flow and muscle force output from the stimulated contractions. These exercise induced increases in vascular resistance usually appear after the first minute of a sustained handgrip contraction (Eklund et al., 1974; Kilbom & Persson, 1982; Saito et al., 1990; Jacobsen et al., 1994). In the present study, force output was linearly associated with blood pressure across the physiological range (Figure 4.7) and did not plateau or decrease at higher blood pressures, indicating that the handgrip contraction was not strong enough to increase leg vascular resistance (Seals, 1989b). The possibility that factors other than the rise in blood pressure contributed to the increase in force when voluntarily contracting the distant muscle was explored by Wright et al. (2000); these could include autonomic influences on the vascular bed or through the -adrenergic receptors on muscle fibres (Martin et al., 1989). However, when the stimulated muscle was elevated to offset the increase in perfusion pressure, no independent influence on force output was demonstrated.

A limitation. Local arterial pressure was used to indicate the relative changes in muscle perfusion pressure as net perfusion pressure, arterial minus venous pressure, could not be measured. The extent to which venous pressure varied during elevating and lowering the legs is uncertain. With the leg muscles contracting rhythmically and the valve system functioning, hydrostatic effects will produce smaller changes in local venous pressures than in local arterial pressures through the action of the active muscle pump (Pollack & Wood, 1949; Folkow et al., 1971; Laughlin, 1987). The other difficulty in interpreting the results is that it is not certain to what extent measurable arterial-venous pressure represents the driving force perfusing the muscle vascular bed given the shunting of arteriovenous anastomoses and, particularly in the legs, the connections between the parallel deep and superficial venous systems. Overall, however, it is undoubtedly true that the changes in net perfusion pressure produced by raising or lowering the legs are somewhat less than indicated by the local arterial pressures measured here. This does not alter the conclusions of the posture-dependent effects but will alter the “gain” of the force-perfusion relationship at the level of muscle physiology. MUSCLE PERFORMANCE 89

Pressor responses to muscle contraction Stimulating the muscle electrically bypassed the central pressor response. As no sizeable cardiovascular responses were observed, it appears that there was also no significant exercise pressor reflex from the working muscle. With the isometric contraction, it can be expected that the muscle mechanoreflex would not evoke a response (Stebbins et al., 1988; Fisher et al., 2005). The workload used in these contractions was similar to that of the leg muscles during standing and not enough to occlude muscle blood flow by raised intramuscular pressure (Barcroft & Millen, 1939; Lind & McNicol, 1967b), explaining why a significant metaboreflex elevation in blood pressure is not observed. There is, however, a suggestion of divergence of the central mean arterial pressures between muscle perfusion conditions (Figure 4.4), with central blood pressure higher during stimulation with low muscle perfusion pressure. This suggests that this level of contraction is near threshold to evoke a metaboreflex response from these muscles.

As the effect of the exercise pressor reflex is minor, the cardiovascular response observed during voluntary contractions originates from the central pressor response. As seen with the stimulated contractions (Figure 4.4), a constant level of drive to the muscle results in a declining force output, even at these relatively low workloads. Thus, voluntary contraction at a constant force output requires a steadily increasing motor command and would be expected to result in a steady increase in blood pressure as seen in the results of Wright et al. (1999). To avert this, subjects tracked the extrapolation of the declining force trajectory and, as this required a constant motor output, blood pressure remained at a constant level. Thus, the changes in central blood pressure when raising and lowering the legs here reflect the true central pressor response to the change in motor command (Goodwin et al., 1972)needed to offset the perfusion-related change in contractility.

With the legs below heart level, muscle perfusion pressure increases and augments muscle contractility so that the same muscle force can be maintained with a smaller voluntary motor command. The lack of regulatory processes that reduces muscle perfusion pressure to match activity benefits the leg muscles during standing where upright posture creates a positive hydrostatic pressure head of more than 80 mmHg (Figure 4.2). This more than sufficient increase in muscle perfusion creates a MUSCLE PERFORMANCE 90

perfusion advantage, whereby the muscle metaboreflex is buffered as muscle contractility is maintained until higher workloads. This means that the exercise pressor reflex plays only a minor role in the regulation of blood pressure in human standing.

Conclusions This study shows that force output from a postural muscle of the leg is highly sensitive to variations in perfusion pressure. This perfusion dependent change in muscle contractility occurs over a short time frame and can be explained by the ionic shifts that mediated changes in membrane potential. These effects are reversible and suggest that a steady-state equilibrium exists between force output and perfusion pressure. The relationship between muscle contractility and perfusion is consistent across muscle groups and produces proportionately similar effects for passive or active changes in perfusion pressure. 91

Chapter 5

Cardiovascular control during standing

Human standing is maintained by continuous contractions of the postural muscles at relatively low workloads. The leg muscles are activated at 15 – 20% of their maximal force output during standing and at these low to moderate workloads, cardiovascular responses are dominated by the central pressor response rather than through peripheral reflexes that originate in the contracting muscles. The central pressor response is described for volitional muscle contraction and a descending motor command from the cerebral cortex. However, as shown in Chapter 3, the dominant motoneuronal drive controlling standing is from a sub- cortical balance system with only a small cortical contribution. Here I ask whether muscle drive from the balance system evokes a central pressor response as it does through the cortical system. Normal subjects performed several balance and non-balance tasks while arterial pressure and heart rate were measured. For comparison, subjects performed a task that was equivalent to standing in which they balanced an inverted mechanical pendulum. This equivalent task excludes descending drive from the balance system. Cardiovascular responses were also measured during voluntary isometric contractions of the calf muscles that were matched to the actual and perceived forces experienced during standing. Testing perceived force contractions in this way indicates the magnitude of the cortically- driven pressor response during standing. The results show that arterial pressure does not change from resting values at the commencement of the contraction used to stand, but increases during the equivalent pendulum task. Maintaining an isometric contraction at the level of the force perceived during standing did not increase arterial blood pressure, whereas muscle contractions matched to the actual force of standing produced a rise in arterial blood pressure. These results show that drive from the balance system does not contribute to the central pressor response during normal standing and that any contribution from the cortical descending drive is below threshold to generate a significant change in blood pressure.

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Introduction

The upright human posture, which we can maintain for extended periods of time, requires sustained activity of the calf muscles to stop the body from toppling. As shown in the previous chapter, the calf muscles depend on the high perfusion pressure developed during orthostatic posture and reducing that pressure causes a decline in contractility. During prolonged standing, fatigue of the leg muscles can be restored partially by increasing perfusion pressure.

Normally when a muscle is worked, the central pressor response (Goodwin et al., 1972) and the muscle pressor reflex (Coote et al., 1971; McCloskey & Mitchell, 1972) raise arterial blood pressure and offset the time-related loss of muscle contractility (Wright et al., 2000). It is estimated that the pressor response, halves the fatigue rate of a muscle worked by volitional contraction (Wright et al., 1999). In the previous chapter it was noted that at workloads similar to those of the calf muscles during standing, the central pressor response accounted for the increase in arterial blood pressure while the muscle pressor reflex made a minor contribution at best.

The experiments of Chapter 3 show that there are two independent sources of neural drive to the posturally active calf muscles during standing; a small motor drive of cortical origin and a large motor drive originating from a presumably sub-cortical balance system to which the vestibular system makes some contribution. While we can expect the cortical drive component to generate a central pressor response (Goodwin et al., 1972; Gandevia & Hobbs, 1990; Winchester et al., 2000), we do not know whether the sub-cortical balance drive feeds the cardiovascular centres that generate the pressor response. The fact that this independent balance system does not contribute to perceptions of force exerted during muscular contraction raises the possibility that it similarly does not contribute to cardiovascular control during postural activity.

These experiments investigate the contribution of the balance system to the regulation of arterial blood pressure during standing. Cardiovascular responses are observed during normal standing and compared to the responses during volitional isometric contractions performed at the actual and perceived levels of force generated by the calf muscles during standing. As shown in Chapter 3, balancing a matched inverted pendulum excludes a contribution to the motor drive from the sub-cortical CARDIOVASCULAR CONTROL 93

balance system and forces reliance on cortical drive. This allows the pressor response during normal standing to be compared with the response that would be generated by cortical drive only at the same workload and in the same orthostatic posture. Methods Experiments were conducted on ten healthy subjects (5 females), aged between 22 and 53 years (mean 31.1). Each had participated in the experiments of the previous chapter that examined their perception of force during standing.

Setup Subjects were studied in seated and upright positions (Figure 5.1). When seated, the legs were horizontal with the knees straight and both feet were secured firmly against a force plate that fixed the ankles at 90°. In three upright postures, subjects: (i) stood freely on a force platform, (ii) were supported in the standing posture by strapping to a rigid post behind them, and (iii) were supported in the standing posture and used the leg muscles to control the inverted mechanical pendulum that simulated the load of the body during standing (described in Chapter 3).



Figure 5.1 Setup A. Subjects stood freely while blood pressure (BP) and heart rate (HR) were measured by finger plethysmography. Torque exerted on the floor was measured from the force plate and EMG activity from the leg muscles was monitored. B. A back support (behind the subject in A) was moved to support the subject in the posture of standing and allowed the leg muscles to be relaxed. Volitional isometric plantarflexion contractions were then made against the force plate to match the actual torque of standing (a) and the perceived torque (p). C. An inverted pendulum matched to the load of the body was balanced to exclude a contribution from vestibular balance drive and force reliance on cortical drive (Chapter 3). D. While seated, subjects produced a plantarflexion contraction matching the torque exerted during standing. CARDIOVASCULAR CONTROL 94

Ankle torque generated by plantarflexion was measured from the force plate and load cells attached to the foot plates of the pendulum. Electromyographic (EMG) activities were recorded and bandpass filtered (30 – 1000 Hz) from surface electrodes placed 6 cm apart longitudinally over the right soleus and tibialis anterior muscles. Beat-to-beat arterial blood pressure (BP) and heart rate (HR) were measured continuously by finger plethysmography (Portapres, Finapres Medical Systems, Amsterdam, The Netherlands). Data were sampled to computer at 2 kHz.

Protocol Subjects refrained from alcohol, caffeine and strenuous exercise for at least 4 hours before testing. Cardiovascular responses were measured during five tasks that were randomised with an intervening rest period of 15 minutes and no more than two tested on a single day. Each task was preceded by a 10-minute seated rest period to ensure stable baseline conditions. BP and HR were measured continuously for 5 minutes with the subjects at rest and then for 5 minutes as they performed the task.

Standing. Testing began with the subject strapped to the support post in the normal standing posture but with the leg muscles relaxed (as in Figure 5.1B). Feedback of EMG activity was provided to ensure complete relaxation. After 5 minutes at rest, the straps and support were removed to start free standing with the eyes open (Figure 5.1A).

Isometric upright. The subject was supported in the normal standing posture with the leg muscles relaxed for 5 minutes before starting a voluntarily contraction of the calf muscles against a force platform to a target force shown on an oscilloscope (Figure 5.1B). In different tests, the target was set as either the actual force exerted or the perceived force exerted by the subject during normal standing, as determined beforehand during Study 1 of Chapter 3. Briefly, the actual force exerted was determined as the subject stood on top of a force platform for 10 s, and the perceived force was determined with the subject supported upright by a rigid post and contracting the calf muscles to match the force experienced during standing.

Inverted pendulum. To begin, the subject was supported in the normal standing posture with the leg muscles relaxed while secured to the support post. After 5 minutes at rest, a lock on the inverted pendulum was released so that the subject began to balance it by contraction of the calf muscles (Figure 5.1C). The pendulum had been matched to the CARDIOVASCULAR CONTROL 95

subject’s body load in advance. Subjects held the pendulum, without seeing it, at a position that required the same ankle torque that had previously been measured for the subject during free standing.

Isometric seated. Testing began with the subject seated and relaxed for 5 minutes. An isometric plantarflexion contraction of the calf muscles then commenced to a target force, shown on an oscilloscope, that was the actual active force exerted during standing (Figure 5.1D). The force of a maximum voluntary plantarflexion contraction was recorded in this posture at a different time.

Analysis Mean arterial pressure was calculated as diastolic plus one-third pulse pressure. Central arterial pressure was estimated by subtracting the hydrostatic pressure of a column of blood, between the heart and the hand, from the arterial pressure recorded in the middle finger of the right hand. Data were averaged every 30 s and normalised to their respective values immediately before the onset of muscle contraction.

The significance of differences in arterial pressure and heart rate at corresponding time points before and during muscle activity were determined by 2-way ANOVA. Repeated-measures ANOVA with post-hoc Bonferroni correction for multiple comparisons were used to determine time effects on blood pressure and heart rate during rest and activity periods. Significance was set at P = 0.05 and group data are presented as means with 95% confidence intervals. Results

Records for a typical subject and group-averaged responses are shown in Figures 5.2 and 5.3 respectively. Arterial pressure and heart rate were stable during the rest period of all conditions. When upright, average resting central arterial pressure was 85.8 mmHg [85.5, 86.1] and average resting heart rate was 83.9 bpm [82.5, 85.3]. When seated, average resting central arterial pressure was 74.6 mmHg [74.4, 74.8] and average resting heart rate was 72.1 bpm [71.8, 72.5]. In only the inverted-pendulum task, arterial pressure rose transiently at the beginning of the baseline rest period by 3.7 mmHg [3.6, 3.8] without a change in heart rate (see Figure 5.3B). Although not statistically significant from the remaining baseline after Bonferroni correction CARDIOVASCULAR CONTROL 96

(P = 0.12), it is noted as only this condition was associated with a change in heart rate on starting muscle contraction.



Figure 5.2 Cardiovascular responses of one subject Data recorded for the five test conditions in one subject show soleus muscle activity (EMG), ankle torque, arterial pressure (BP), and heart rate (HR) with the subject at rest during the first half of the record and then during a sustained plantarflexion contraction of the calf muscles. An increase in BP is observed when subjects balanced the body-matched pendulum (B) and when they matched the actual force generated during standing while seated (C) or supported upright (D). These responses are not seen at the commencement of free standing (A) or when matching the perceived force of standing (E).

Average plantarflexion force exerted by the calf muscles during standing was 18.2% [17.8, 18.6] of the measured maximal voluntary contraction force. When subjects matched the actual force generated during standing, in both the upright and seated positions, arterial pressure increased. Illustrated for one subject in Figure 5.2, both systolic and diastolic pressure increased during the first minute after the onset of muscle activity to reach a new stable level for the remainder of the contraction. Similar rises in systolic and diastolic pressure were observed at the commencement of balancing the mechanical pendulum. In contrast, blood pressure did not increase when the subject CARDIOVASCULAR CONTROL 97

contracted the calf muscles to stand or when the subject matched the perceived force of contraction during standing.

The same was observed in the group responses (Figure 5.3). There were no significant changes from resting values in central arterial pressure or heart rate during the 5 minutes of quiet standing (Figure 5.3A). When subjects made an isometric contraction to match the perceived force experienced during standing, equivalent to 5.8% [5.6, 5.9] of maximum voluntary force, there were likewise no significant changes in arterial pressure or heart rate. However, when subjects made a voluntary isometric contraction to match the actual force of standing, central arterial pressure increased to a maximum of 5.8 mmHg [5.7, 5.9] or 6.8% (P < 0.001 by ANOVA) after 3 minutes (Figure 5.3B). A comparable rise in arterial pressure was observed when subjects contracted the leg muscles to balance the inverted pendulum, when central arterial pressure increased by 6.2 mmHg [6.1, 6.3] or 7.2% (P < 0.001 by ANOVA) after 3 minutes.



Figure 5.3 Group data for upright tests A. Mean arterial pressure and heart rate were recorded as subjects were supported upright and at rest for the first 5 minutes of each record (group means ± S.E.M.). At time T0, subjects began the contraction of the calf muscles that continued for the second 5 minutes of each record. They either stood freely (black) or were supported upright and matched the force perceived during standing (blue). B. In the same format, relative changes in arterial pressure and heart rate are displayed for the inverted pendulum task (yellow) and when subjects matched the actual force exerted during standing (red). Subjects were supported upright throughout both conditions. Data from the normal, free standing condition (black) are superimposed for comparison. CARDIOVASCULAR CONTROL 98

The increases in arterial pressure when matching the actual force of standing and balancing the inverted pendulum were accompanied by changes in heart rate. For the matched isometric contraction the effect was smaller and transient as heart rate decreased by 3.1 bpm [3.0, 3.2] or 3.7% after 60 s (P < 0.05 by ANOVA) before recovering towards resting values during the remaining 240 s of muscle contractions. When balancing the inverted pendulum, heart rate decreased at the onset of muscle activity by a maximum of 5.9 bpm [5.8, 6.0] or 7.0% after 90 s (P < 0.001 by ANOVA). Heart rate recovered slightly during the contraction but remained significantly below resting values (e.g. 4.6 bpm [4.5, 4.7] or 5.5% at 180 s).

When subjects made the same voluntary contraction while seated, matching the actual force output of standing, arterial pressure rose by 6.5 mmHg [6.4, 6.6] or 8.7% after 180 s (P < 0.001 by ANOVA; Figure 5.4), similar to the 6.8% rise measured for the same contraction when upright. In contrast to the upright postures, heart rate increased at the onset of muscle activity and remained elevated for the duration of the contractions (e.g. increased by 4.0 bpm [3.9, 4.1] or 2.6% after 180 s (P < 0.001 by ANOVA).



Figure 5.4 Group data for seated and upright contractions Relative changes in arterial pressure and heart rate are shown as subjects matched the actual force experienced during standing in the seated (green) and upright (red) postures (group means ± S.E.M.). Records are in the format of Figure 5.3 with 5 minutes of rest followed by

5 minutes of contraction starting at T0. Data of the upright condition are those of Figure 5.3B.

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Discussion

During normal standing, activation of the postural muscles does not increase arterial pressure above resting levels. In contrast, the same level of contraction driven by the voluntary cortical system, whether to produce an isometric contraction or to balance a load equivalent to the body, increases arterial pressure significantly. These results show that the control of muscle activity by volitional and automatic balance-related processes produces distinctly different cardiovascular responses because motor drive from the automatic balance system does not evoke a central pressor response (Goodwin et al., 1972).

Neither arterial pressure nor heart rate changed when subjects went from the supported upright position to active standing. The same was observed when they remained supported upright but commenced a voluntary contraction of the calf muscles at the perceived level of force during standing, which was approximately one-third of the actual contraction force of standing. If, as indicated by the studies of Chapter 3, we assume that the size of the descending cortical commands were the same for the free- standing and perceived-force contractions, we can also conclude that any central pressor response has little effect on arterial blood pressure at this low level of drive. The perceived-force contractions were only 6% of maximal force output for the calf muscles. At similar contraction forces, Fallentin et al. (1985) showed that sustained low-level contractions (7% MVC) of the elbow flexor or extensor muscles did not produce a cardiovascular response. Thus, the absence of any contraction-related cardiovascular response during standing is consistent with the small cortical motor drive and the absence of a pressor response from the automatic balance drive.

Characteristic central pressor responses were observed when seated subjects matched the actual force of standing by contracting the calf muscles isometrically. Arterial pressure reached a new steady-state during the contraction over the time scale (60 – 90 s) of the perfusion-pressure effects shown in Chapter 4, suggesting that a new perfusion-contractility equilibrium had been reached. When subjects were immobilised in the upright position and matched the actual force of standing, a similar pressor response was observed. This similarity between the two responses indicates that the central pressor response operates with equal gain with both postures and the CARDIOVASCULAR CONTROL 100

contractility advantage that the leg muscles gain from the high hydrostatic perfusion pressure does not necessarily remove the need for, or effect of, a pressor response to maintain contractility during a sustained contraction. In other words, muscle performance could improve if there was a pressor response.

We should consider that these studies focus only on replicating the activity of the calf muscles during standing. There are many more postural muscles active during standing compared with the inverted pendulum or isometric force tasks. However, although the muscles that stabilise the trunk and head during standing add to the total mass of exercised muscles, there was no evidence of a metaboreflex response when subjects went from the supported posture at rest to free standing. Similarly, a contribution from the central pressor response would most likely be trivial since these postural muscles are presumably also activated by the sub-cortical balance system (Ali et al., 2003).

In contrast to real standing, balancing the body-equivalent mechanical pendulum produced an increase in arterial blood pressure. Even though a significant metaboreflex is unlikely at these workloads (Chapter 4), this difference in the pressor response appears to be central in origin as the calf muscles contracted at equivalent workloads. As proposed by Fitzpatrick et al. (1992), Chapter 3 shows that balancing the pendulum excludes the vestibular sub-cortical motor drive to the postural muscles and forces a reliance on cortical motor drive. Thus I conclude here that motor drive from the sub- cortical balance system is not involved in the regulation of arterial blood pressure during normal standing.

The voluntary contractions of the calf muscles while upright produced unexpected changes in heart rate. Associated with the increases in blood pressure, a modest, transient slowing of heart rate was observed when subjects matched the actual force of standing and this was more pronounced and prolonged when they balanced the mechanical pendulum. This did not happen in the seated posture, where contraction increased pulse rate. Generally, heart rate increases during muscle activation due to the withdrawal of cardiac vagal tone by cortical motor commands (Rowell & O'Leary, 1990; DiCarlo & Bishop, 1992) and, perhaps, by reflex activity from muscle mechanoreceptors (Stebbins et al., 1988; Hayes & Kaufman, 2001; Gladwell & Coote, CARDIOVASCULAR CONTROL 101

2002). The explanation for this cardiac slowing is not clear. The two different postures will bring about differences in lower-body venous volume, central venous pressure, sympathetic tone, and the set point of the baroreceptor reflex. Considering that both resting arterial pressure and heart rate were lower in the seated posture, any combination of these factors could contribute to cardiac slowing during muscle activation in an upright posture. Further studies are required to understand this unexpected behaviour.

Significance for human standing The dependence on a hydrostatically induced change in local perfusion pressure coupled with the absence of a pressor response suggest that, during standing, central blood pressure regulation is dominated by arterial and cardiopulmonary baroreceptors. Support for this comes from studies that show that arterial blood pressure and heart rate in upright posture, after passive head up tilt or from actively standing up, are indistinguishable after 60 seconds (Tanaka et al., 1996; Heldt et al., 2003). As muscle fatigue develops during a continuous volitional contraction, the increasing motor drive that must be delivered to compensate for the lost force output increases arterial pressure through the central pressor response (Wright et al., 1999). The inability to raise arterial pressure during standing by a contribution from the central pressor response and a sub- threshold contraction to evoke a muscle pressor reflex could be a factor contributing to orthostatic hypotension. The voluntary leg muscle contractions we make if we experience pre-syncopal symptoms could work because they generate the pressor response in addition to increasing venous return.

Conclusion This study reveals a differential effect on the cardiovascular system between the automatic balance task of standing and an equivalent voluntary contraction in the same orthostatic posture. These differences are a consequence of dual sources of motoneuronal drive for postural control. The larger neural drive from the balance system does not produce a cardiovascular response while the smaller cortical drive is likely to be sub-threshold to evoke a pressor response during normal standing.

102

Chapter 6

Conclusions and speculations

Contracting muscles rely on the coordination of neuromuscular and cardiovascular control processes, which bear many similarities. Both engage forward anticipatory mechanisms based on centrally-generated signals associated with the motor command, and both rely on feedback from peripheral signals originating in contracting muscles. Through these similarities, this thesis set out to explore the neuromuscular and cardiovascular interactions in human standing, an activity critical to our species and demanding for both control processes.

The results of these experiments provide interesting novel concepts on the origin of the neural drive to the leg muscles during standing and its implications for cardiovascular control. The cortical processes that control voluntary motor tasks are overshadowed by a strong output from a separate motor system for balance tasks. Similarly, the rise in blood pressure that typically accompanies and is essential for sustained muscular activity is absent during standing because of these dual neural control processes.

The work of this thesis covers a very broad range of physiological processes and has answered some questions but poses many more. Any aspect could be considered in great depth but in this brief discussion I will focus on aspects of the control of human standing that were particularly interesting. Posture and balance

A range of psychophysical studies explored the neural processes that control human standing. These revealed two sources of motor drive that appear to act independently. Volitional motor actions are accompanied by a conscious perception of the consequence of that action. When we lift an object, its weight and movement are fused with the perception of the action. However, this process is not entirely straightforward for CONCLUSIONS 103

standing when the “lifting” is the muscle contraction that stops the body falling forward. Subjects only perceived a small fraction of the force exerted by the calf muscles during standing. The results of Chapter 2 provide the explanation. Figure 6.1 outlines the conclusions of Chapter 2 and Chapter 3 in a scheme of the sensorimotor neural processes that subserve human standing.

Figure 6.1 Schema of sensorimotor control of standing Two separate motor systems contribute to the force exerted during standing with each motor system generating its own distinct afferent information. Muscular contraction generated by volitional motor drive from the cortex (blue) is associated with reafference signals (orange) that when compared with the corollary discharge gives rise to the perception of self-generated force. The reafference signal may be composed of Golgi tendon organ (GTO) and cutaneous afferents as well as increased muscle spindle afferent signals due to the co-activation of alpha (-MN) and gamma (-MN) motoneurons. When afferent signals associated with muscle contraction that is driven by the balance system is compared to the corollary discharge it is interpreted as external stimuli (exafference, grey) and does not contribute to the sense of force exerted during standing.

As seen in the two deafferented subjects studied, sensations related to the force exerted by muscles during volitional contractions arise in the absence of afferent sensory inflow. In normal subjects also, illusory sensations of movement and resistance can result from volitional contraction in the absence of afferent inflow (Walsh et al., CONCLUSIONS 104

2010). With afferent inflow available, however, perceptions are very different. As proposed by von Holst & Mittlestaedt (1950) a central efference copy of the motor action allows the total afferent signal to be separated into components that are the result of the action (reafference) and components imposed by external stimuli (exafference). In Figure 6.1, these are illustrated as the two aspects of perception.

It is unclear how efference-related or corollary discharge signals within the central nervous system “contribute” to the perceptions of the outcome of volitional contractions. Action and outcome form a unified sensation that also embodies properties of the load or external contact point. It is clear from the experiments of Chapter 2 that what has been considered to be a centrally-generated corollary discharge signal could actually represent a reafference signal from peripheral receptors. An unusual aspect of previous interpretations of the proprioceptive senses has been that movement and position have been attributed to peripheral signals arising in muscle, skin and joint receptors whereas force exerted by the muscles has been attributed, largely, to centrally generated signals (McCloskey, 1981b; Gandevia, 1996). In light of the findings here, this view should be reconsidered.

When standing, the cortex does not issue most of the descending drive to the muscles controlling balance. Thus, there is no efference copy that allows reafference to be extracted. Reafference of motor drive issued by the sub-cortical balance system appear to the cortical perceptual system as exafference (grey connections in Figure 6.1). Only the smaller cortical command for standing results in a reafferent signal that feeds perception, resulting in only a weak sense of force exerted.

Whether the balance system also operates through a principle of reafference, separating its own actions from the consequences of external balance perturbations is uncertain. It is tempting to suppose that it does because the principle of reafference is ubiquitous to many biological systems, accounting for the complex decoding of visual inflow to generate a stable internal representation of the external visual world down to the neural processes of simple primitive sensory arcs (von Helmholtz, 1867; von Holst & Mittelstaedt, 1950; Bell, 1989). It could be argued that there is no need for this type of processing in a system that has the sole role of resisting gravito-inertial forces. The evidence for it comes from the results of vestibular stimulation experiments, which CONCLUSIONS 105

show that balance reflexes are distributed to different muscles according to their ability to resist gravity and their relative orientation with the sensor organs (Britton et al., 1993; Fitzpatrick et al., 1994a; Cathers et al., 2005). It is conceivable then that at this sub-conscious level, an efference copy of the balance system motor command is used to process all somatosensory afference to identify effective muscle groups through their reafference. Cardiovascular regulation during standing

How does the central nervous system coordinate the cardiovascular actions associated with the two separate motor systems during standing balance? Well, it appears that it doesn’t. Figure 6.2 outlines the conclusions of Chapter 4 and Chapter 5 in a scheme of the cardiovascular control processes that act during standing and volitional contractions. There was no cardiovascular response when the body went from a resting, upright position to balancing the body in standing (Chapter 5). This lack of a cardiovascular pressor response can be directly attributed to the neural processes that oversee balance control. Cardiovascular centres in the brainstem receives inputs from both central signals that are associated with the cortical motor command and from the periphery.

During standing, cortically-generated motor commands were not strong enough to stimulate a cardiovascular pressor response as they represent only a small portion of the total motor drive to the muscles. A cortically-driven command that would also replace the command from the balance system does evoke a pressor response so we can conclude that motor activity driven by the balance system does not evoke a pressor response.

It seems odd that the balance system, with its dominant role in the motor control of upright standing, is not involved in cardiovascular regulation during standing. However, it appears that the sub-cortical balance system is concerned with cardiovascular regulation but that it operates differently, being more concerned with blood distribution and maintaining caudal blood pressure by regulating sympathetic outflow than perfusing active postural muscles (Yates, 1992; Yates & Miller, 1994). There is well documented anatomical and neurophysiological evidence for pathways for vestibular influences on medullary cardiovascular centres (Balaban & Beryozkin, 1994; Yates et al., 1994; Porter & Balaban, 1997; Stocker et al., 1997; Balaban & Porter, CONCLUSIONS 106

1998). In human subjects, who should depend more on cardiovascular homeostasis to combat gravitational forces, a vestibulo-sympathetic reflex has been demonstrated during brief head flexion (Shortt & Ray, 1997; Ray & Hume, 1998), off-vertical head and body rotation manoeuvres (Hume & Ray, 1999; Kaufmann et al., 2002), and through galvanic mastoidal stimulation (Bent et al., 2006).

Figure 6.2 Cardiovascular regulation Added to the scheme on sensorimotor control in Figure 6.1 is a scheme of the cardiovascular control processes that act during standing and volitional contractions. The pressor response, activated in a feedforward manner by the cortical motor command (solid purple line), is not activated by the motor drive originating from the balance system. The balance system exerts its own influence on the cardiovascular system by influencing medullary centres and sympathetic drive that redistribute blood to offset gravitational influences (broken purple line).

More questions and more work

With the expansive approach and scope of this thesis, significant observations made at each stage have not been pursued in detail. As with any scientific pursuit, answering one question poses many more. Further investigations are required and planned to understand some of the results obtained. A selection of these are discussed here.

CONCLUSIONS 107

Balance, posture and movement

 Consider this person making a volitional contraction of the postural muscles while supported and making an equivalent contraction to stand. The studies of Chapter 3 show that simply by releasing the straps and having the body exposed to the consequences of the gravitational field engages entirely different neural control processes that ultimately appear to have a similar influence on the motoneuronal drive to the muscles.

Perhaps the most interesting questions arising from this observation are: “What turns on the balance system?” and “Why?” The results of these and previous studies (Fitzpatrick et al., 1994b) indicate that standing can be maintained without undue effort or difficulty by the cortical motor system.

Performing a concurrent cognitive task while standing results in some deterioration in cognitive performance (Kerr et al., 1985; Lajoie et al., 1993). This is viewed as evidence that standing is “attentionally demanding” and requires cortical resources. However, the results of Chapter 3 indicate a not-overwhelming role for the cortex in standing balance considering that loads such as the matched inverted pendulum used in these studies can be balanced with ease. It could be suggested that the cortex undertakes more of a monitoring function during balance control, assessing the motor output of the balance system and intervening only when necessary, such as during large or anticipated postural disturbances. It remains a puzzle that the cortex maintains a perceptual independence and treats actions of the balance system as exafference, as if they originate from the external world. This of course is how vestibular symptoms are reported: “It was as if I was being pushed over.”

We might assume that having this independent system specifically for balancing the body upright in gravity, distinct from balancing other loads, confers some CONCLUSIONS 108

significant advantage. However, it might be that evolution did not reach the solution in this simplistic way. Geotropism is a primitive and fundamental biological process. In the animal kingdom the balance sense is the most primitive, appearing before the visual and auditory senses and certainly before the evolution of the cerebral cortex, which we believe is responsible for our form of conscious perception. The balance system provides tonic activation of the extensor muscles to oppose gravity since its sensory organs are specifically sensitive to gravitational acceleration. A way to view this dual system is that the cortical system, which is concerned primarily with movement, has evolved on top of this automatic balance system, trusting it to perform its balance duties and incorporating its predictable actions into a forward model of movement planning.

Although the upright body during standing can be seen as a simple inverted- pendulum load that can be maintained by controlling ankle posture (Fitzpatrick et al., 1992; Loram et al., 2001), it is clear from the studies presented here that balancing the body and controlling the posture or alignment of a joint involve completely different, and largely isolated, neural control processes. The cortex has its own solution to postural control and it is important to make the distinction with balance control. One view of cortical postural control, the “equilibrium point theory” is that movement is simply the transition between neurally coded postures (Bizzi et al., 1992; Feldman & Levin, 1995). Another view is that posture and movement are organised separately. Within the cortex, there is evidence that different neuronal populations are active during postural tasks and voluntary movement tasks (Kurtzer et al., 2005; Scott, 2008) and these differences are reflected in human postural control of the limbs (Chew et al., 2008). However, the studies of this thesis suggest that neither of these views are appropriate to understand the motor control of balance, which has its own agenda related to gravitational alignment rather than creating body postures.

Several questions on the operation of the balance system are currently being addressed. First, what turns it on? The hypothesis being tested is that the balance system does indeed act in accord with the reafference principle and is activated if gravitationally relevant reafference is detected. To test this, subjects will balance the inverted pendulum (Figure 3.5) while a galvanic current, modulated according to the sway of the pendulum, stimulates the vestibular system to create a signal correlated with the leg muscle activity; i.e. mimicking a reafference into the balance system. The CONCLUSIONS 109

appearance of vestibular reflexes and a reduced perception of force applied will be taken as evidence that the balance system has been engaged.

Is standing analagous to the respiratory motor control where both cortical and sub-cortical controls exist and the cortex can assume control as needed? Does the cortex “release” control to the balance system and can it resume it during balance? To answer this, studies are underway to determine if corticomuscular coherence is modulated in a task-dependent way during standing. When subjects make a voluntary contraction of the calf muscles equivalent to that of standing, corticomuscular coherence is high compared with that during standing (Figure 3.7), probably reflecting the reduced cortical drive and reafference during standing (Baker, 2007). Exploiting this, experiments are underway to determine whether this coherence shows a task dependence when standing in different conditions, indicating that the cortex can regain or yield control to the balance system.

Muscle contractility The muscle contractility studies of Chapter 4 establish that a steady-state equilibirum exists between muscle perfusion and muscle contractility. Essentially, the motor function of a muscle can be considered as a family of length-velocity-tension curves or surfaces that are linked by perfusion. The reversible effects of perfusion are profound in relation to these mechanical factors and equally significant for motor control. This reversible equilibrium is not the generally accepted view of muscle action and muscle fatigue, which is generally thought of as a one-way process that takes considerable time and rest for recovery. When reduced perfusion pressure that created a decline in force output was restored, force output recovered to the predicted decline in force output had the legs received the higher perfusion pressure for the duration of the experiment (Figure 4.4). A first hypothesis would be that reducing muscle perfusion for a period would exacerbate the longer-term fatigue process and contractility would not return to the original trajectory on restoring perfusion but stabilise on a lower trajectory. The simple addition of the two processes and absence of a cross-term in the double exponention decline in force output (see Figure 4.8) suggests that the processes underlying the long-term fatigue are buffered against rapid changes in perfusion. The method used here of controlling the perfusion-dependent contractility process, which appears to be explained by interstitial K+ concentration, should help identify other factors contributing to muscle fatigue. CONCLUSIONS 110

Without a recognised exercise pressor response during standing the leg muscles rely on local properties in the lower limb to support the on-going muscle activity required to keep the body upright and balanced. There is a large perfusion advantage of the head of hydrostatic pressure created simply by being upright. The resulting increase in muscle pefusion leads directly to an increase in muscle contractility (Chapter 4). The nature of the muscle contraction and load could be particularly significant for the efficiency of the leg muscles during standing. Here, I have investigated muscle function using isometric contractions. However, during standing, the legs contract with greater variation because of the unstable nature of the load relative to the stiffness of the ankles that can be exerted through balance reflexes (Fitzpatrick et al., 1992), and in a pulsatile manner because the passive stiffness of the ankles is less than the load stiffness of the body (Loram & Lakie, 2002; Fitzpatrick, 2003). Pulsatile movement could improve muscle perfusion through periods of low intramuscular pressure and through the muscle pump, which depends on a full venous system. The cardiovascular significance of these load-dependent contraction patterns are being investigated using protocols developed in these studies.

Cardiovascular regulation during standing In Chapter 5, when subjects balanced a mechanical inverted pendulum while their body was restrained, blood pressure increased while heart rate fell significantly whereas when performing an isometric contraction of equivalent strength while seated or supported upright, both blood pressure and heart rate increased. This suggests a fundamental difference in cardiovascular regulation between these two volitional motor activities, which from the results of Chapter 3 (see Figures 3.7 and 3.8) are unlikely to be related to the involvement of the balance system. In an early study, Asmussen et al. (1939), referred to in a review paper by Smith & Porth (1991), observed that heart rate decreased when subjects voluntarily contracted the leg muscles while tilted 50° head- up. It could be speculated that these differences are related to postural factors and blood volume but vestibular involvement in a cardiovascular regulatory role is possible (Yates, 1992; Yates & Miller, 1994). On the other hand, reflex and late postural responses to perturbations applied to the arm when balancing upright pendula are very different from the responses when making an equivalent isometric contraction (Chew, 2009). This suggests that the two motor activities could generate different patterns of CONCLUSIONS 111

reafference, which is shown to be the basis of the sense of effort and thus linked to the central pressor response. Again, a combination of psychophysical and cardiovascular studies are planned to answer these questions.

…. and on this thesis

The investigations of this thesis took a tortuous path and are essentially presented in reverse chronological order. Beginning with an investigation of the significance of the cardiovascular responses to the leg muscles during standing, described in Chapter 5, it was necessary to backtrack to understand the effects of perfusion pressure on the postural muscles described in Chapter 4. Understanding the central origin of the motor command driving the pressor response led to the investigation of perceived force and effort during standing, described in Chapter 3. In turn, it was soon apparent that it was necessary to backtrack to understand the general process of force perception. This study, described in Chapter 2, reshapes our ideas of the neural basis of the perception of our actions and provides a solution to a century-long debate of the peripheral versus central origin of force sensation. Not until the end of these studies did their meaning become clear.

112

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