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Does a decrease in seat height modify the effect of cadence on activation of the triceps surae during.. 2008

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  DOES A DECREASE IN SEAT HEIGHT MODIFY THE EFFECT OF CADENCE ON ACTIVATION OF THE TRICEPS SURAE DURING CYCLING?    by   RYAN PETER CAWSEY   B.H.K., The University of British Columbia, 2003      A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF   MASTER OF SCIENCE   in    THE FACULTY OF GRADUATE STUDIES  (Human Kinetics)    THE UNIVERSITY OF BRITISH COLUMBIA  (Vancouver)     August 2008   © Ryan Peter Cawsey, 2008    ii  Abstract   Introduction:  Several authors have demonstrated that, while cycling at a constant power output, EMG activity from the gastrocnemius increases systematically with increases in pedaling cadence, but that soleus EMG remains unchanged (Marsh & Martin 1995; Sanderson et al. 2006). The reason for this differential effect of cadence on the muscles of the triceps surae is unclear. Whatever factor(s) are responsible, it is assumed that, as they vary, the differential electromyographic response will vary accordingly. Decreasing the seat height has been shown to alter the kinematic characteristics of cycling (Too, 1990). The first objective of this study was to examine the effect of a decrease in seat height on the kinematics and muscle activation of the lower limb. The second objective was to investigate the effect of seat height on the relationship between cadence and triceps surae activation and, in doing so, to reveal possible factors mediating the response to changes in cadence.  Methods:  Participants pedaled a cycle ergometer at 200 Watts for five minutes at each of five cadences (50, 65, 80, 95, 110 rpm) and at each of two seat heights (100% and 90% trochanteric height). Kinematics of the lower limb were calculated from digitized video records of reflective markers placed on the skin over seven bony landmarks. EMG data were collected from eight lower-limb muscles.  Results:  The most notable findings were 1) that activation of the gastrocnemii was less in the low-seat condition and, contrary to what the findings of past research would suggest, was not associated with changes in muscle length; 2) that the medial and lateral gastrocnemii responded differently to changes in cadence at each seat height, suggesting that the functional roles of these muscles in cycling differ; 3) that several factors, including muscle length, muscle velocity, ankle angle and the direction of muscle action, were not responsible for the differential effect of cadence on activation of the soleus and gastrocnemius. Future research should investigate afferent feedback from proprioceptors in the knee joint and knee extensor muscles as possible factors mediating the effect. iii  Table of Contents   Abstract..................................................................................................................................................... ii  Table of Contents.................................................................................................................................. iii  List of Figures......................................................................................................................................... vi  Acknowledgements............................................................................................................................... ix 1.0 Review of Literature..................................................................................................................... 10  1.1 Length-Tension Relationship in Skeletal Muscle.................................................. 10   1.1.1 Intrinsic Characteristics of Skeletal Muscle Fibres.............................................. 10   1.1.2 Interaction of Intact Muscle with the Nervous System........................................ 13  1.2 Force-Velocity Relationship in Skeletal Muscle..................................................... 16   1.2.1 Concentric Muscle Action....................................................................................... 16   1.2.2 Eccentric Muscle Action.......................................................................................... 17  1.3 Relationship between Muscle Force, Velocity and Activation......................... 19   1.3.1 Differences in Global Muscle Activity between Concentric and Eccentric Actions.............................................................................................. 19   1.3.2 Motor Unit Recruitment during Eccentric Muscle Action.................................. 21   1.3.3 Differential Activation of the Triceps Surae......................................................... 22  1.4 Structural Differences between the Soleus and Gastrocnemius Muscles.................................................................................................... 23   1.4.1 Pennation Angle....................................................................................................... 23   1.4.2 Fibre Type................................................................................................................ 23  1.5 EMG Activity of the Triceps Surae during Cycling............................................. 24   1.5.1 Typical Activity Patterns........................................................................................ 24   1.5.2 Selection of Preferred Cadence.............................................................................. 25   1.5.3 Effect of Seat Height................................................................................................ 25   1.5.4 Effect of Cadence..................................................................................................... 26 2.0 Statement of the Problem.......................................................................................................... 27 3.0 Methods.................................................................................................................................................. 28  3.1 Data Collection..................................................................................................................... 28  3.2 Data Processing................................................................................................................... 30 iv   3.3 Statistical Analysis.............................................................................................................. 32 4.0 Results...................................................................................................................................................... 32  4.1 Joint Angle 33   4.1.1 Ankle Angle.............................................................................................................. 33   4.1.2 Knee Angle............................................................................................................... 33   4.1.3 Hip Angle………...................................................................................................... 33  4.2 Muscle Length...................................................................................................................... 34   4.2.1 Soleus Length........................................................................................................... 34   4.2.2 Medial Gastrocnemius Length............................................................................... 34   4.2.3 Lateral Gastrocnemius Length……....................................................................... 35   4.2.4 Tibialis Anterior Length......................................................................................... 35   4.2.5 Semimembranosus Length...................................................................................... 35   4.2.6 Vastus Lateralis Length.......................................................................................... 36   4.2.7 Rectus Femoris Length............................................................................................ 36   4.2.8 Gluteus Maximus Length........................................................................................ 36  4.3 Muscle Velocity.................................................................................................................... 37  4.4 EMG......................................................................................................................................... 37   4.4.1 Soleus EMG.............................................................................................................. 37   4.4.2 Medial Gastrocnemius EMG.................................................................................. 38   4.4.3 Lateral Gastrocnemius EMG................................................................................. 38   4.4.4 Tibialis Anterior EMG............................................................................................ 39   4.4.5 Semimembranosus EMG........................................................................................ 40   4.4.6 Vastus Lateralis EMG............................................................................................. 40   4.4.7 Rectus Femoris EMG.............................................................................................. 40   4.4.8 Gluteus Maximus EMG.......................................................................................... 41 5.0 Discussion.............................................................................................................................................. 41  5.1 Objective I: Effect of Seat Height................................................................................ 41  5.2 Objective II: Effect of Cadence..................................................................................... 46  5.3 Objective III: Combined Manipulations................................................................... 47   5.3.1 Muscle Length.......................................................................................................... 47   5.3.2 Muscle Velocity........................................................................................................ 48   5.3.3 Joint Angle................................................................................................................ 48   5.3.4 Direction of Muscle Action..................................................................................... 49  5.4 Additional Factors.............................................................................................................. 49   5.4.1 Pennation Angle....................................................................................................... 49   5.4.2 Moment Arm Length............................................................................................... 49   5.4.3 Fibre Type................................................................................................................ 50  5.5 Differential Functioning of the Gastrocnemii......................................................... 50  5.6 Agonist-Antagonist Coactivation................................................................................. 51  5.7 Muscle Length Calculations........................................................................................... 52  5.8 Limitations............................................................................................................................. 53 6.0 Conclusion............................................................................................................................................ 53 7.0 References............................................................................................................................................. 55  Appendix 1: Figures from Grand Means.................................................................................... 64  Appendix 2: Figures from Individual Participants................................................................. 88  Appendix 3: Behavioural Research Ethics Board Certificate of Approval.................. 122  Appendix 4: Participant Consent Form....................................................................................... 124    v  vi  List of Figures  Appendix 1: Grand Means Figure 1: Mean ankle joint angle and angular velocity for each cadence and seat-height condition.......................... 65 Figure 2: Mean knee joint angle and angular velocity for each cadence and seat-height condition........................... 66 Figure 3:  Mean hip joint angle and angular velocity for each cadence and seat-height condition.............................. 67 Figure 4: Soleus length, velocity and EMG for each cadence and seat-height condition............................................ 68 Figure 5: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition.................. 69 Figure 6: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition.................. 70 Figure 7: Tibialis anterior length, velocity and EMG for each cadence and seat-height condition............................ 71 Figure 8: Semimembranosus length, velocity and EMG for each cadence and seat-height condition....................... 72 Figure 9: Vastus lateralis length, velocity and EMG for each cadence and seat-height condition............................. 73 Figure 10: Rectus femoris length, velocity and EMG for each cadence and seat-height condition............................ 74 Figure 11: Gluteus maximus length, velocity and EMG for each cadence and seat-height condition........................ 75 Figure 12: Soleus EMG for each seat-height and cadence condition.......................................................................... 76 Figure 13: Medial gastrocnemius EMG for each seat-height and cadence condition................................................. 77 Figure 14: Lateral gastrocnemius EMG for each seat-height and cadence condition................................................. 78 Figure 15: Lateral gastrocnemius EMG integral for each cadence and seat-height condition.................................... 79 Figure 16: Tibialis anterior EMG for each seat-height and cadence condition........................................................... 80 Figure 17: Tibialis anterior EMG integral for each cadence and seat-height condition.............................................. 81 Figure 18: Semimembranosus EMG for each seat-height and cadence condition...................................................... 82 Figure 19: Vastus lateralis EMG for each seat-height and cadence condition............................................................ 83 Figure 20: Rectus femoris EMG for each seat-height and cadence condition............................................................ 84 Figure 21: Gluteus maximus EMG for each seat-height and cadence condition.......................................................... 85 vii  Figure 22:  The EMG integral value for the low seat condition subtracted from the EMG integral for the high seat condition for each seat height condition....................................................................................................................... 86 Figure 23:  Soleus, medial gastrocnemius and lateral gastrocnemius maximum shortening velocity for each cadence and seat-height condition................................................................................................................................ 87  Appendix 2: Individual Participants Figure 24: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 1............ 89 Figure 25: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 1.................................................................................................................................................................. 90 Figure 26: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 1.................................................................................................................................................................. 91 Figure 27: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 2............ 92 Figure 28: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 2.................................................................................................................................................................. 93 Figure 29: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 2.................................................................................................................................................................. 94 Figure 30: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 3............ 95 Figure 31: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 3.................................................................................................................................................................. 96 Figure 32: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 3.................................................................................................................................................................. 97 Figure 33: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 4............ 98 Figure 34: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 4.................................................................................................................................................................. 99 Figure 35: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 4.................................................................................................................................................................. 100 Figure 36: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 5............ 101 Figure 37: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 5.................................................................................................................................................................. 102 Figure 38: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 5.................................................................................................................................................................. 103 Figure 39: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 6............ 104 Figure 40: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 6.................................................................................................................................................................. 105 Figure 41: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 6.................................................................................................................................................................. 106 Figure 42: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 7............ 107 Figure 43: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 7.................................................................................................................................................................. 108 Figure 44: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 7.................................................................................................................................................................. 109 Figure 45: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 8............ 110 Figure 46: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 8.................................................................................................................................................................. 111 Figure 47: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 8.................................................................................................................................................................. 112 Figure 48: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 9............ 113 Figure 49: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 9.................................................................................................................................................................. 114 Figure 50: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 9.................................................................................................................................................................. 115 Figure 51: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 10.......... 116 Figure 52: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 10................................................................................................................................................................ 117 Figure 53: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 10................................................................................................................................................................ 118 Figure 54: Soleus length, velocity and EMG for each cadence and seat-height condition from Participant 11.......... 119 Figure 55: Medial gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 11................................................................................................................................................................ 120 Figure 56: Lateral gastrocnemius length, velocity and EMG for each cadence and seat-height condition from Participant 11................................................................................................................................................................ 121  viii  ix  Acknowledgements   I would like to express my appreciation to my friends, family, colleagues and committee members without whose guidance and support this project would not have been possible. To Dr. Tim Inglis, Dr. Romeo Chua and Dr. Mark Carpenter for their time spent in the development of this and preliminary research projects; to my supervisor, Dr. David Sanderson, whose encouragement during my undergraduate degree led me to pursue graduate studies, and whose enthusiasm for teaching was unwavering throughout my Masters degree; to my colleagues, Scott, Julia, Karine and Lexi with whom it was a pleasure to spend many days in the lab; to my parents for their unquestioning support; and to Jen whose love and friendship guided me through many challenges, my sincere thanks.  10  1.0 Review of Literature 1.1 Length-Tension Relationship in Skeletal Muscle  The capacity of intact muscle to produce tensile force is dependent on the muscle’s length. This relationship is brought about by the intrinsic characteristics of isolated skeletal muscle fibres, as well as through the interaction of intact muscle and the nervous system. These two mechanisms are discussed in the following sections.  1.1.1 Intrinsic Characteristics of Skeletal Muscle Fibres  Seminal experiments on isolated animal skeletal muscle fibres demonstrated that when a muscle fibre is tetanically stimulated, an optimal sarcomere length exists for the production of tension; at greater or lesser sarcomere lengths, the tension is less than maximal (Ramsey & Street, 1940; Gordon et al., 1964; Edman, 1966; Gordon et al., 1966). These findings were extended by Rack and Westbury (1969) who examined sarcomere lengths in the intact cat soleus muscle throughout a range of physiological joint angles and demonstrated that the sarcomere length-tension relationship of isolated muscle fibres is present in an intact joint-muscle-tendon unit. Subsequent studies on skeletal muscle in humans (Walker & Schrodt, 1973; Lieber et al., 1994; Maganaris, 2001, , 2003) have similarly shown that the range of sarcomere lengths seen within the physiological range of operating muscle lengths falls on a portion of the sarcomere length-tension curve that affects the tension-producing capacity of intact skeletal muscle throughout its physiological range of lengths.  The sarcomere length-tension relationship of an isolated muscle fibre can be explained by the sliding filament theory, a model of muscle contraction proposed simultaneously and independently by A.F. Huxley and R. Niedergerke (1954) and by H.E. Huxley and J. Hanson (1954) which states that changes in length of the sarcomere are brought about by the thin, actin filament sliding past the thick, myosin filament. It was later discovered that the sliding movement is caused by the incremental pulling action of the heads of the myosin filaments on the actin filaments at binding sites called cross-bridges (Huxley, 1957a; Huxley, 1969; Huxley & Simmons, 1971). This mechanism for the development of muscle force was supported by the finding that the isometric tension in a tetanic muscle fibre is directly proportional to the number of cross-bridges that are formed (Huxley, 1957a; Huxley, 1957b; Gordon et al., 1966). The length of an isolated muscle affects the muscle’s force-producing capabilities by changing the amount of overlap of the actin and myosin filaments within the sarcomere and, thus, changing the number of cross-bridges that are formed. A sarcomere is at its optimal length for force production when the myofilaments overlap such that 11  the most cross-bridges are formed. This has been shown to occur in the muscles of the lower limb in humans at sarcomeres lengths of 2.5 µm to 2.8 µm (Walker & Schrodt, 1973). When a muscle is stretched beyond its optimal length, the myofilaments within the sarcomere are pulled longitudinally resulting in less overlap and the formation of fewer cross-bridges. At a length of 4.3 µm, the sarcomere is not able to produce any force because there is no overlap of the actin and myosin filaments. Likewise, when a muscle is drawn shorter than its optimal length, the actin filaments overlap with one another, creating fewer binding sites for the formation of cross-bridges. At a length of 1.25 µm, the actin filaments overlap to such a degree that the sarcomere cannot generate any force. Several confounding factors make it difficult to test the length-tension relationship of skeletal muscle in vivo. The first is that the moment about a given joint is the net product of all active synergist and antagonist muscle forces and their moment arm lengths; the second is that moment arm lengths bear non-linear relationships to joint angle (Kawakami et al., 1998; Maganaris et al., 1998a); and the third is that muscle fibre pennation angle, which is known to affect the force-producing capacity of muscle (see section 1.4.1), also bears a non-linear relationship to joint angle (Maganaris et al., 1998b; Maganaris & Baltzopoulos, 1999). To complicate matters further, muscle moment arm lengths and pennation angle measurements at rest or from cadaveric specimens differ from measurements made from active muscle (Narici et al., 1996; Kawakami et al., 1998; Maganaris et al., 1998a, 1998b; Maganaris & Baltzopoulos, 1999; Maganaris et al., 1999), and may therefore result in erroneous conclusions when studying in vivo muscle function. Despite these difficulties, many attempts have been made to identify the portion of the length-tension curve along which skeletal muscle operates in vivo (see table 1). The findings of these studies have revealed considerable variability in the length-tension properties of skeletal muscle between different species of animals and between different muscles within the same species. Some of this variability may be attributed to differences in the structure of myofilaments between species. The length of thick myosin filaments is reported as being 1.6 µm and has been found to vary little among vertebrates (Page & Huxley, 1963; Walker & Schrodt, 1973). On the other hand, the length of thin actin filaments has been found to be 0.9-1.0 µm in frogs (Page & Huxley, 1963; Sonnenblick & Skelton, 1974), 1.0-1.1 µm in rodents and monkeys (Walker & Schrodt, 1973; Robinson & Winegrad, 1979) and 1.2 µm in humans (Walker & Schrodt, 1973). The sarcomere length-tension relationship is influenced by the length of the actin filaments: a longer actin filament will lead to a longer ascending limb and a rightward shift of the plateau region to longer sarcomere lengths (Walker & Schrodt, 1973; Loren et al., 1996). Many of the studies that have attempted to identify the portion of the length-tension curve along which skeletal muscle operates in vivo used the 12  original sarcomere length-tension curve described by Gordon et al. (1966) based on frog specimens. Comparing these data to those of other animals would lead to erroneous conclusions and may be one explanation for the inter- animal variability in the reporting of the length-tension properties of skeletal muscle.  A second possible explanation is that the portion of the length-tension curve along which the same muscle from the same species operates in vivo may change to reflect the muscle’s specific pattern of use, as Herzog et al. (1991a) suggested in their cross-sectional study comparing competitive cyclists to runners. A similar hypothesis was postulated by Lieber et al. (1994) concerning differences in the length-tension relationship between the human extensor carpi radialis brevis, a muscle generally used for the skilled manipulation of objects, and that of the large muscles involved in bipedal locomotion, such as rectus femoris.  Table 1. Studies that have investigated the portion of the length-tension curve at which skeletal muscle operates in vivo. Reference Animal Muscle Portion of the Length-Tension Curve (Herzog et al., 1991b) Human Gastrocnemius Ascending Limb  (Maganaris, 2003) Human Gastrocnemius Ascending Limb  (Maganaris, 2001) Human Soleus and Tibialis anterior Ascending Limb and Plateau  (Rack & Westbury, 1969) Cat Soleus Ascending Limb  (Leedham & Dowling, 1995) Human Biceps brachii Ascending Limb  (Lieber et al., 1992) Fish (Kryptopterus bicirri) Muscle(s) dorsal to the vertebral column  Plateau  (Lieber & Brown, 1992) Frog 7 muscles of the lower limb Descending  (Rome & Sosnicki, 1991) Carp (Cyprinus carpio) Muscle(s) dorsal to the vertebral column Ascending Limb and Plateau  (Lieber & Boakes, 1988) Frog (Rana pipiens) Semitendinosus Descending Limb and Plateau  (Mai & Lieber, 1990) Frog (Rana pipiens) Semitendinosus Descending Limb  (Lieber et al., 1994) Human Extensor carpi radialis brevis Descending Limb and Plateau  (Cutts, 1988) Human Cadavers 9 muscles of the lower limb Ascending Limb and Plateau  (Nordstrom & Yemm, 1974) Rat Masseter Descending Limb  (Weijs & van der Wielen-Drent, 1982) Rabbit (Oryctolagus cuniculus) Masseter Descending Limb  (Hertzberg et al., 1980) Rabbit Masseter, Digastric Ascending Limb  (Ichinose et al., 1997) Human Vastus Lateralis All regions    13  1.1.2 Interaction of Intact Muscle with the Nervous System  In addition to the mechanical properties of the sarcomere, changes in the way that the nervous system evokes muscular contraction also contribute to the effect of muscle length on the tension produced. Rack and Westbury (1969) described changes in a muscle’s twitch profile with changes in the muscle’s length. They did so by electrically stimulating the nerve supply of the intact soleus muscle in anaesthetized cats while recording the resulting isometric tension produced by the muscle. It was found that, at short muscle lengths, there was a decrease in the peak twitch tension, the rate of tension increase, and the half relaxation time. Similar changes had been found previously by directly stimulating isolated preparations of the sartorius muscle in frogs (Doi, 1921; Hartree & Hill, 1921; Jewell & Wilkie, 1960) and were later demonstrated in humans by electrically stimulating the skin surface above the nerve supply of the tibialis anterior, abductor digiti minimi, biceps brachii and brachialis muscles (Gandevia & McKenzie, 1988). Rack and Westbury (1969) also found that as muscle length decreased, the firing frequency required to maintain a given tension increased, a finding that was later replicated in the human abductor digiti minimi muscle (Gandevia & McKenzie, 1988). This increase in firing frequency is in agreement with the previously described changes in twitch profile with decreased muscle length: twitches with a smaller peak tension, a slower rise to peak and a faster decrease from peak tension would require a higher firing frequency to achieve the same degree of summation of individual twitch contractions with a continuous train of electrical impulses (Zajac & Young, 1980). To keep tension constant when a muscle is in a shortened position, the nervous system must increase the stimulus frequency.  The results of experiments that have examined the effect of muscle length on the stimulus rate required to maintain a given level of muscle force have been supported by subsequent studies in which single motor unit action potentials were recorded with in-dwelling electrodes during isometric contractions at various muscle lengths. Miles et al. (1986) recorded single motor unit action potentials from the masseter muscle at various jaw angles while participants bit down onto a customized strain gauge. Participants were asked to begin biting with the smallest possible force and to increase slowly until the onset of the motor unit from which EMG activity was being recorded. It was found that the recruitment threshold for motor unit activity increased with increasing length of the masseter muscle. Vander Linden et al. (1991) recorded single motor unit action potentials from the tibialis anterior muscle while participants produced a constant magnitude of dorsiflexor torque with their foot strapped into a strain gauge- instrumented orthosis at three different ankle angles. They found that the rate of motor unit firing for a given magnitude of dorsiflexor torque was greater when the tibialis anterior muscle was at short lengths. Using a similar 14  design, Pasquet et al. (2005) recorded single motor units action potentials from the tibialis anterior while participants performed maximal voluntary isometric contractions at 10° of plantarflexion and 10° of dorsiflexion. Their findings agreed with those of Vander Linden et al. (1991): when the ankle was in a position of dorsiflexion, tibialis anterior fascicle length was shorter, the torque produced during maximal voluntary contraction was reduced, mean surface EMG activity increased, motor unit recruitment threshold decreased, and motor unit discharge rate increased. Christova et al. (1998) recorded single motor unit action potentials during isometric twitch contractions from the biceps brachii muscle at three different elbow angles and found that motor unit firing frequency was higher at short muscle lengths. Each of these findings demonstrates a change in motor unit activation by the nervous system to accommodate changes in sarcomere length.  Interestingly, the effect of muscle length on the recruitment threshold and firing frequency of motor units in the medial gastrocnemius muscle is opposite to that seen in muscles previously investigated. Kennedy and Cresswell (2001) recorded single motor unit action potentials from the medial gastrocnemius muscle and surface EMG activity from the soleus muscle while participants gradually increased the magnitude of plantarflexor torque with their foot strapped to an instrumented footplate. Recordings were made once with the knee completely extended and a second time with the knee flexed to ninety degrees; this allowed for the length of the biarticular gastrocnemius muscle to be manipulated while the monoarticular soleus muscle remained at a fixed length. It was found that, with the knee in a flexed position, the onset of motor unit activity from the medial gastrocnemius occurred at a greater magnitude of plantarflexor torque and that the frequency of action potentials at onset was higher. The functional significance of this relationship lies in the articular nature of the gastrocnemius and soleus muscles.  The medial and lateral heads of the gastrocnemius muscle attach superior to the knee joint at the medial and lateral condyles of the femur, respectively. The soleus muscle attaches inferior to the knee joint at the posterior- superior aspect of the tibia and fibula. All three muscles share a distal attachment at the posterior aspect of the calcaneus via the Achilles tendon (Seeley et al., 2006). Given these attachment sites, the gastrocnemius acts both at the knee and the ankle, while the soleus acts only at the ankle. Flexion of the knee causes a shortening of the gastrocnemius but has no effect on the length of soleus. These two muscles present a unique situation in which the requirement to produce a given plantarflexor torque can be shared to optimize the length-tension properties of each muscle. As the knee moves from a position of full extension to one of full flexion, the gastrocnemius operates on a less effective portion of its length-tension curve. Steindler (1977) described a situation referred to as ‘active insufficiency’ in which biarticular muscles shorten to such an extent that they cannot produce any amount of tension. 15  This phenomenon has been cited by several authors as a possible reason for the changes in the recruitment of the gastrocnemius and soleus with changes in knee angle (Herzog & ter Keurs, 1988; Herzog, 2000; Kennedy & Cresswell, 2001; Arampatzis et al., 2006). The hypothesis of derecruitment due to “active insuffiency” is founded on the overarching principle that human movement is optimized for minimum energy expenditure (Zarrugh et al., 1974; Sparrow & Irizarry-Lopez, 1987; Sparrow & Newell, 1994; Bertram & Ruina, 2001). Following this principle, Arampatzis et al. (2006) concluded that, since muscle activation requires metabolic energy, the motor control system reduces muscular activity to maximize the economy of force generation. When the gastrocnemius is actively insufficient, its contribution to plantarflexor torque would be minimal, even at near maximal levels of activation; therefore, it would be a metabolically prudent measure to decrease the recruitment of this muscle and increase the recruitment of a plantarflexor synergist, soleus, that is at a more effective position along its length-tension curve. Muscles that have been shown to exhibit an increase in recruitment with a decrease in length cannot share their workload with a mono-articular synergist of different length-tension properties (Miles et al., 1986; Vander Linden et al., 1991; Christova et al., 1998). In these muscles, there is no option but to increase recruitment as the muscle moves to a less effective length in order to maintain a constant joint torque.  Tamaki et al. (1997) proposed several possible mechanisms through which the decrease in gastrocnemius activation at short muscle lengths takes place. The first is a decrease in cortical activation of the motor neurons innervating the medial gastrocnemius. This mechanism is made less likely, though not impossible, by the increase in activation of the soleus muscle that has been shown to occur concurrently with the decrease in recruitment of medial gastrocnemius motor units (Kennedy & Cresswell, 2001). The second proposed mechanism is an inhibition or disfacilitation of gastrocnemius α-motoneurons through peripheral afferent input regarding muscle length from spindle receptors and regarding joint position from cutaneous and joint receptors. The third mechanism is a heteronymous excitatory influence on soleus α-motoneurons from proprioceptors in the knee extensor muscles.  Surface EMG recordings offer a global perspective of the summed activity from individual motor units and can allow for different conclusions to be drawn than those from an intramuscular EMG signal. However, changes in surface EMG activity that occur with changes in muscle length are subject to the confounding effect of muscle fibre movement with respect to the recording electrode. Never-the-less, experiments that have examined the relationship between muscle length and the surface electromyogram in the muscles of the triceps surae have yielded results that support the findings of the studies using intramuscular EMG described previously. Sale et al. (1982) recorded surface EMG activity from the entire triceps surae muscle group using large (20 x 15 cm) electrodes placed over 16  both the gastrocnemius and soleus muscles and found that the electrical activity for a given plantarflexor torque was less when the ankle was in a position of plantarflexion.  Cresswell et al. (1995) recorded surface EMG activity from the soleus and the gastrocnemius muscles during voluntarily generated isometric plantarflexor torques of various relative magnitudes throughout a range of knee angles with the ankle angle held constant. The amplitude of EMG activity from gastrocnemius was found to decrease with increasing knee flexion, whereas the amplitude of EMG activity from soleus remained unchanged. This effect on gastrocnemius activity was independent of the magnitude of plantarflexor torque. Identical changes in the relative activity of the soleus and gastrocnemius during maximal voluntary isometric contractions were found by Miaki et al. (1999) and Shinohara et al. (2006).  Investigating the dynamic action of the triceps surae, Tamaki et al. (1997) recorded surface EMG from the gastrocnemius and soleus muscles while participants performed plantarflexor movements at three different ankle angular velocities and with the knee held fixed at three different angles. It was found that EMG activity from gastrocnemius decreased with increasing knee flexion, while EMG activity from soleus remained unchanged or decreased. This trend was independent of the angular velocity of plantarflexion.  As the gastrocnemius muscle is shortened, there is a greater amount of overlap between myofilaments and a greater number of cross-bridges would be expected to form under the EMG electrode fixed to the surface of the skin. The result of this would be an increase in the EMG signal, independent of any changes to the recruitment of the muscle. That the experiments by Sale et al. (1982) and Cresswell et al. (1995) yielded decreases in surface EMG with muscle shortening despite the possibility of there being an increase in the number of cross-bridges under the EMG electrode, supports the notion that these differences were brought about by changes in muscle recruitment. 1.2 Force-Velocity Relationship in Skeletal Muscle  In addition to being influenced by muscle length, a muscle’s force-producing capacity is dependent on the velocity with which the muscle is shortening or lengthening. This force-velocity relationship is different for shortening versus lengthening muscle actions, as well as for isolated versus intact muscle.  1.2.1 Concentric Muscle Action  The force-velocity relationship during muscle shortening in animals was first demonstrated by maximally electrically stimulating isolated muscle preparations (Levin & Wyman, 1927; Fenn & Marsh, 1935; Hill, 1938; Katz, 17  1939). From a series of early experiments on the frog sartorius muscle, Hill (1938) developed the now well-known equation relating muscle tension to the velocity of shortening:  (T + α) · (v + β) = (T0 + α) β  where T is the external load v is the shortening velocity T0 is the maximal isometric load α and β are constants  This equation states that when an isolated muscle is electrically stimulated and is free to shorten, the velocity with which it does so decreases in a hyperbolic fashion, as the load imposed on the muscle increases. With a continued increase in load, the velocity continues to decrease until the shortening velocity reaches zero and the muscle action becomes isometric. This relationship was later found to be true for individual motor units and muscle fibres (Edman, 1979).  The concentric torque-velocity relationship in humans was first demonstrated by recording elbow flexor torque with a fly-wheel dynamometer during voluntary activation of the intact elbow flexor muscles (Hansen & Lindhard, 1923; Wilkie, 1950). This relationship was found to be similar in shape to the concentric force-velocity curve of isolated muscle. More recently, isokinetic, velocity-controlled dynamometers were used to describe the torque-velocity relationship of the human knee extensor muscles and, again, yielded a curve similar in shape to that found in isolated muscle (Hislop & Perrine, 1967; Thorstensson et al., 1976; Evetovich et al., 1997).  The determination of the force-velocity relationship of intact muscle bears many of the same difficulties as those inherent in determining the length-tension relationship of intact muscle. Due to the non-linearity of changes in moment-arm length with changes in joint angle, a given rate of change of joint angular displacement is not equivalent to the rate of change of muscle length. Also, in examining muscle function in vivo, it is difficult to control for the confounding effect of changes in motor unit recruitment by the nervous system with changes in the velocity of muscle action.  1.2.2 Eccentric Muscle Action  The force-velocity curve during eccentric actions of isolated muscle exhibits a continuation of the curve seen in concentric actions: as the load applied to an isolated muscle fibre increases beyond the maximal load able to be withheld in an isometric contraction, the lengthening velocity of the fibre increases in a hyperbolic fashion (Hill, 1938). The eccentric curve contains two separate regions that are more distinctive than in the concentric curve. In 18  the first region, force increases up to 1.5 times the maximal isometric force with modest increases in lengthening velocity. In the second region, a plateau is seen in which force remains constant with additional increases in lengthening velocity (Katz, 1939; Joyce et al., 1969; Flitney & Hirst, 1978; Lombardi & Piazzesi, 1990).  The increases in the speed of lengthening with increases in load during maximal eccentric action occur at a slightly lower rate than the decreases in shortening speed with increases in load in maximal concentric action. Two mechanisms are thought to create this effect. 1) While, in maximal concentric actions, cross-bridge detachment is a chemical process involving the break-down of ATP, in maximal eccentric actions, cross-bridges are broken mechanically (Flitney & Hirst, 1978). The mechanical cross-bridge detachment may occur at a higher muscle tension than the chemical cross-bridge detachment. 2) The time required for cross-bridge detachment in eccentric actions is longer than in concentric actions (Lombardi & Piazzesi, 1990). This would allow a greater time for single action potential summation and would lead to greater muscle tension.  The force-velocity relationship of skeletal muscle in situ with voluntary activation differs from that of isolated preparations. Several researchers have investigated the torque-velocity relationship of the knee extensor muscles by measuring knee extensor torque during voluntary maximal concentric, isometric and eccentric muscle actions at various knee joint angular velocities using a velocity-controlled dynamometer (Perrine & Edgerton, 1978; Westing et al., 1988). Concentric knee extensor torque was found to decrease with increasing angular velocity, in agreement with the force-velocity curve of isolated muscle fibres. Eccentric knee extensor torque, however, did not differ from isometric knee extensor torque with increasing angular velocity. Perrine & Edgerton (1978) postulated that an inhibitory neural mechanism may have limited muscle tension during eccentric actions to avoid muscle injury.  Dudley et al. (1990) improved upon the study of Westing et al. (1988) by including experimental conditions in which the knee extensor muscles were maximally activated via transcutaneous tetanic electrical stimulation. In the voluntary muscle activation conditions, results were similar to those reported by Westing et al. (1988). However, with transcutaneous electrical stimulation, the torque-velocity relationship for both concentric and eccentric actions resembled the force-velocity curve seen in isolated muscle fibres. This provides support for the hypothesis that the differences in the torque-velocity relationship of skeletal muscle in vivo and the force-velocity relationship of isolated muscle fibres is brought about by tension-limiting changes in the activation of intact muscle by the central nervous system. 19   Seger and Thortensson (2000) compared the torque-velocity relationship of the knee extensor muscles during maximal voluntary and submaximal transcutaneously stimulated muscle actions. Their findings were in agreement with those seen previously: for the maximal voluntary efforts, concentric torque decreased as velocity increased, while eccentric torque remained equal to the maximal isometric torque with further increases in velocity. In the submaximal conditions, concentric torque again decreased as the velocity of shortening increased; however, eccentric torque continued to increase with further increases in the velocity of lengthening. This finding further supported the notion that the decrease in knee extensor torque during eccentric muscle action is due to a protective down-regulation of knee extensor activity. Aagard et al. (2000) demonstrated that with eccentric resistance training, the protective neural inhibition can be overcome.  Several possibilities exist for the origin of the inhibitory, or disfacilitory, influence of muscle lengthening on force generation. These include 1) a decrease in cortical drive to the motor neuron pool and 2) a reduction in α- motoneuron excitability brought about through homonymous negative feedback from Golgi tendon organs, muscle spindles and cutaneous and joint receptors (Seger & Thorstensson, 2000).  The length-tension properties of skeletal muscle influence the force-velocity relationship. Since an optimal length exists for the production of force, as the length of a muscle approaches the optimal value, there is a greater velocity of shortening for a given applied load (Bahler, 1968). Bahler et al. (1968) combined the two-dimensional length-tension and force-velocity curves to express the relationship between length, tension and velocity with a single three-dimensional plot. 1.3 Relationship between Muscle Force, Velocity and Activation  A controlled, voluntary, eccentric muscle action differs in many ways from an eccentric action that occurs despite a maximal volitional effort to oppose muscle lengthening. From a simplistic perspective, a controlled eccentric action can be brought about by decreasing the level of muscle activation until the muscle torque is less than the load torque. However, eccentric actions have been found to rely upon specific strategies of muscle activation not seen with concentric actions. The differences in muscle activation between concentric and eccentric actions will be discussed in the following sections.  1.3.1 Differences in Global Muscle Activity between Concentric and Eccentric Actions  Hill (1938) first described the relationship between the work done by an isolated muscle and the heat that the muscle produces. He demonstrated that, for a given load, the heat produced during a concentric action was 20  greater than that produced during an isometric action. This difference in heat was found to increase linearly as the load applied to the isolated muscle was increased; that is, the heat produced by the muscle rose as the velocity of muscle shortening increased. For eccentric actions, the heat produced was found to be less than that produced during isometric actions. This followed the same linear relationship as seen in concentric actions: as the speed of muscle lengthening increased, the heat produced decreased.  Abbot and Bigland (1953) tested Hill’s findings in intact human muscle by measuring the oxygen consumption of participants riding a bicycle ergometer. In some experimental conditions, the pedals of the ergometer were driven backward by a motor while participants resisted the motion with a constant forward load, the magnitude of which was indicated through a visual feedback display. Pedaling without the motor led to a generalized concentric action of the muscles of the legs, while pedaling against the force of the motor led to a generalized eccentric muscle action. It was found that, for a given load, oxygen consumption increased with increasing velocity of concentric action but remained constant with increasing velocity of eccentric action.  The amplitude of surface EMG activity during isometric muscle action increases with muscle force (Lippold, 1952; Bigland & Lippold, 1954). This relationship has generally been found to be linear in muscles of uniform fibre composition and non-linear in muscles of mixed fibre composition (Bigland-Ritchie & Woods, 1974). From this relationship, and from what was already known about the heat generation and oxygen consumption of muscle function with changes in the velocity of muscle action, Bigland and Lippold (1954) hypothesized – and subsequently confirmed – that the amplitude of surface EMG activity during submaximal efforts would increase as the speed of muscle shortening increases, but that no change would be seen with increases in the rate of muscle lengthening. Since then, several other studies have found the EMG activity in the knee extensors during maximal voluntary efforts to be lower during eccentric actions, despite a higher torque generation, than in concentric actions (Eloranta & Komi, 1980; Tesch et al., 1990; Westing et al., 1991). This finding parallels the force-velocity relationship of muscle in vivo during eccentric actions discussed previously.  Twitch interpolation provides another method of assessing the degree of voluntary motor unit activation in a muscle or muscle group. For this technique, a supramaximal electrical stimulus is delivered to the muscle during a maximal voluntary contraction while force or torque output is recorded. The force or torque output during the initial voluntary effort is compared to that seen with the supramaximal stimulation. Since the degree of motor unit activation is proportional to force or torque output, the difference between the two measurements gives an indication 21  of the degree of voluntary motor unit activation relative to the maximal activation achieved through artificial stimulation.  Several studies have used twitch interpolation technique to examine the relationship between the velocity of action and voluntary activation in intact human muscle. When participants were asked to generate maximal knee extensor torques against a velocity-controlled dynamometer, the degree of voluntary activation was greater during concentric muscle actions, despite a greater absolute magnitude of knee extensor torque during eccentric actions (Westing et al., 1990). Babault et al. (2001) examined a wider range of shortening and lengthening speeds in the knee extensor muscle group and demonstrated nearly identical results. These findings support those of Bigland and Lippold (1954) who used surface EMG activity as a global measure of muscle activity.  Since it is the electrical activity in muscle that leads to the generation of force, the mechanism(s) responsible for the reduction in muscle activation during eccentric actions is likely the same as that responsible for the plateau in the force-velocity relationship of intact muscle during eccentric actions discussed previously.  1.3.2 Motor Unit Recruitment during Eccentric Muscle Action  Henneman’s Size Principle (Henneman, 1957) states that, as the requirement for muscle tension increases, low-threshold, slow-twitch motor units are recruited first, followed by high-threshold fast-twitch motor units, and as the requirement for muscle tension decreases, the fast twitch motor units are derecruited first, followed this time by the slow twitch motor units. This principle has found to be true in isometric and concentric movements, both voluntary and reflexive in nature (Milner-Brown et al., 1973; Yemm, 1977; Binder et al., 1983; Calancie & Bawa, 1985a, 1985b).  Two studies, Nardone et al. (1989) and Howell et al. (1995), reported that the typical order of recruitment is reversed during eccentric muscle actions. They reason that such a reversal would facilitate the execution of a submaximal eccentric muscle action by allowing a more rapid decrease in muscle tension (Nardone et al., 1989). For example, consider a situation in which the elbow flexor muscle group must actively lengthen to lower an applied load. If the motor units of the elbow flexors were derecruited according to the Size Principle, that is, in the order of largest to smallest, the rate of muscle lengthening, and thus elbow extension, would be limited by the relaxation time of the low-threshold, slow-twitch motor units. By derecruiting the slow-twitch fibres first, the speed of the movement can be increased while still being actively controlled by the firing of fast-twitch motor units. 22   In response to the papers by Nardone et al. (1989) and Howell et al. (1995), Bawa and colleagues published two articles in which they argued against a departure from Henneman’s Size Principle during eccentric actions (Bawa & Jones, 1999; Stotz & Bawa, 2001). They explained the results of the group arguing the affirmative in the following way: 1) The apparent selective recruitment of fast twitch motor units during lengthening actions actually occurred during brief, high velocity, concentric actions arising from stretch reflex activation. These reflexive concentric actions, Bawa states, are observable in the plots of torque vs time, and have been shown to recruit motor units according to size (Calancie & Bawa, 1985b). 2) The action potentials that are recorded with an intramuscular needle electrode are from a small range of motor units that are recruited at similar magnitudes of muscle tension. The larger action potentials seen during eccentric actions were, therefore, from motor units close in proximity along the continuum of motor unit sizes to those that were active during concentric actions.  1.3.3 Differential Activation of the Triceps Surae  The relative activity of the soleus and gastrocnemius has been shown to differ with changes in the velocity of muscle action. EMG activity from surface (Nardone & Schieppati, 1988) and intramuscular electrodes (Nardone et al., 1989) during submaximal eccentric actions has been shown to be less in the soleus and greater in the gastrocnemius than during concentric action of the same muscle tension. Similar changes in the relative activity of the soleus and gastrocnemius muscles have been found in cats during rapid paw-shaking movements (Smith et al., 1980). High-frequency flexion-extension movements of the ankle were elicited by placing the cat’s paw in water or by sticking tape to the plantar pads. During these movements, the lateral gastrocnemius, composed primarily of fast- twitch fibres was active while the soleus, composed primarily of slow-twitch fibres, was completely silent. In their research on the activity of the triceps surae muscle group in cats that were allowed to move freely, both Smith et al. (1977) and Walmsley et al. (1978) found no other movement during which the lateral gastrocnemius was active without the soleus; the selective recruitment of fast-twitch ankle extensors could only be induced under the high velocity demands of paw-shaking. The studies by Smith et al. (1980), Nardone and Schieppati (1988) and Nardone et al. (1989) support the notion that the muscles of the triceps surae, and even specific motor units within an individual muscle can be selectively activated to best accommodate the velocity of muscle action. 23   1.4 Structural Differences between the Soleus and Gastrocnemius Muscles 1.4.1 Pennation Angle  Pennation angle is defined as the angle between the orientation of a muscle’s fibres and the muscle’s line of pull (Enoka, 2002). The larger the pennation angle, the less are a muscle`s maximal tension and velocity of shortening (Wickiewicz et al., 1983). Soleus has been found to have a pennation angle of 20°, lateral gastrocnemius 9° and medial gastrocnemius 16° (Wickiewicz et al., 1983; Cutts, 1988).  Pennation angle has been found to change with muscle length and tension. Studies using ultrasonic imaging techniques have found that, as the ankle became more plantarflexed, the pennation angle of the human medial gastrocnemius increased (Narici et al., 1996; Maganaris et al., 1998b). With the ankle held constant in a neutral position, the pennation angle increased as the magnitude of isometric tension increased from rest to the maximal magnitude. These findings suggest that a portion of the length-tension and force-velocity properties of skeletal muscle are due to tension- and velocity- dependent changes in pennation angle.  1.4.2 Fibre Type  Sarcomeres within a muscle lie in both a serial and parallel arrangement. The parallel arrangement is related to the muscle’s cross-sectional area which is proportional to the muscle’s force-producing capacity. The serial arrangement allows for a greater degree of displacement and more rapid muscle shortening (Wickiewicz et al., 1983). The ratio of a muscle’s fibre length to cross-sectional area gives an indication of the functional role of the muscle. Compared to the other major muscles of the leg, the plantarflexors have sarcomere arrangements that favour the production of force at the expense of displacement. When making comparisons within the plantarflexors, the gastrocnemius is best suited for displacement, while the soleus is best suited for the production of force.  Edgerton et al. (1975) used a histochemical staining technique on muscle samples taken from thirty-two cadavers to map the fibre type composition of the soleus and gastrocnemius muscles (see table 2). The soleus was found to be composed of a greater percentage of slow-twitch fibres and a smaller percentage of fast-twitch fibres. Both muscles contained approximately the same proportion of intermediate (fast-oxidative glycolytic) fibres. The medial and lateral gastrocnemii were found to have little difference in their fibre compositions.   24  Table 2: muscle fibre type composition of the soleus and gastrocnemius muscles (Edgerton et al. 1975)  Soleus Gastrocnemius Slow Oxidative: 64% 52% Fast Oxidative Glycolytic: 17% 16% Fast Glycolytic: 19% 31%  1.5 EMG Activity of the Triceps Surae during Cycling  The cycle ergometer provides a useful paradigm for the study of muscle function for several reasons: 1) it allows for the easy averaging of many cycles of the same movement, 2) compared to walking or running, its constrained path of motion reduces the degrees of freedom and leads to a decrease in inter-cycle variability and 3) it controls for the potentially confounding effect of having to bear one’s own weight, while allowing for a movement involving the major muscles of the lower limb. In this section, the typical activity pattern of several muscles of the lower limb will be described and some of specific characteristics of cycling that relate to the study of muscle function will be addressed.  1.5.1 Typical Activity Patterns  In their review of muscle activity during pedalling, Hug and Dorel (2007) provide a comprehensive summary of the typical EMG profile over the crank cycle for each of the major muscles of the lower limb. Several key observations were made from the activation profiles observed in nine separate studies (Ericson, 1986; Jorge & Hull, 1986; Ryan & Gregor, 1992; Hug et al., 2004a; Hug et al., 2004b; Hug et al., 2006a; Hug et al., 2006b; Dorel et al., 2007; Duc et al., 2008). These observations are: 1) the gluteus maximus is active from TDC to approximately 130°, 2) the vastus lateralis becomes active just prior to TDC until a point just prior to BDC, 3) the activity pattern of rectus femoris is similar to that of vastus lateralis, although its activity begins and ends slightly earlier than that of vastus lateralis, 4) the tibialis anterior is most active from 270° to TDC, 5) the medial and lateral gastrocnemii are active from approximately 30° to 270°, opposite to periods of tibialis anterior activity, 6) the soleus is active from TDC to just prior to BDC, and 7) the activity pattern of the semimembranosus differs between studies with some accounts of activity from TDC to BDC, and others from TDC to 270°. Hug and Dorel also noted that the activity patterns of the mono-articular lower-limb muscles tend to be less variable than those of bi-articular muscles.   25  1.5.2 Selection of Preferred Cadence  Whereas the preferred stride frequency for a given speed of weight-bearing locomotion is that at which metabolic cost is minimized (Bertram & Ruina, 2001), the same principle does not apply to the determination of preferred cadence in cycling. The cadence associated with minimum energy expenditure varies linearly as a function of power output and ranges from approximately 40 rpm at 40 Watts to 60 rpm at 330 Watts (Banister & Jackson, 1967; Seabury et al., 1977); however cadences as high as 80 rpm have been found to be optimal in highly trained cyclists (Coast & Welch, 1985). This range of cadences is much lower than that preferred by elite cyclists during endurance training or racing (90-100 rpm) and by noncyclists (70-80 rpm) (Marsh & Martin, 1993).  Seabury et al. (1977) offered the following explanation for the existence of a cadence associated with minimum metabolic cost for a given power output. If power output is kept constant, the average force applied to the pedals needed to turn the crank decreases as cadence increases (Sanderson, 1991). Since a muscle fibre has an upper limit of the force that it can produce, as the force requirements of the entire muscle increase, a greater number of muscle fibres are recruited to generate a greater combined force. With a greater number of muscle fibres being recruited comes an increase in the energy lost through heat. This phenomenon creates the lower limit of the cadence optimized for minimum energy expenditure. Increasing cadence would decrease the muscles’ force requirements; however, at high speeds of action, more energetically inefficient muscle fibres are recruited. Also, at high cadences, there is an increase in the viscous friction between active and non-active muscle fibres, leading to an increase in heat production. These mechanisms make up the upper limit of the cadence optimized for energy expenditure  Several studies have shown through optimization modelling techniques that, although not perfect predictors of preferred cadence, joint moment- and muscle stress-related functions reach minima at pedalling frequencies close to the self-selected value (Redfield & Hull, 1986; Hull & Gonzalez, 1988a, 1988b; Marsh et al., 2000).  1.5.3 Effect of Seat Height  Changes in seat height during cycling lead to changes in limb segment kinematics which lead to changes in muscle lengths, moment arm lengths, and thus joint moments of the lower limbs. The optimal seat height is typically defined as that which is associated with minimum metabolic cost and has been found to be equal to 100% of trochanteric height, measured as the rectilinear distance from the ground to the greater trochanter of the femur with the participant in a standing position (Shennum & deVries, 1976; Nordeen-Snyder, 1977); however, for short 26  periods of high intensity, anaerobic work, Hamley and Thomas (1967) found that a seat height of 102% trochanteric height was optimal.  Amoroso (1994) investigated the effects of changes in seat height on the activation of the muscles of the lower limb during cycling. Participants rode at a power output of 200 W, at a cadence of 80 rpm at three seat heights: preferred height, -10% of preferred and +5% of preferred height. EMG activity from both soleus and gastrocnemius were least in the low-seat condition compared to the two higher seat conditions, with a larger effect being seen in gastrocnemius. The mechanism responsible for this change in muscle activation could be the same as that responsible for the changes in activation of the gastrocnemius with varying degrees of knee angle reported by Sale et al. (1982), Cresswell et al. (1995), Tamaki et al. (1997) and Kennedy and Cresswell (2001); that is, a disfacilitation of the gastrocnemius at short muscle lengths due to its ineffectiveness in generating tension.  1.5.4 Effect of Cadence  Several efforts have been made to quantify the effect of cadence changes on the activity of the triceps surae muscle group during cycling. Ericson et al. (1985) used a bicycle ergometer with a weight-loaded braking system to examine the activity of eight muscles of the lower limb at four different cadences (40, 60, 80, 100 rpm) at a constant power output of 120 Watts. They found that EMG activity of the medial gastrocnemius increased with increases in cadence, while EMG activity of the soleus remained mainly unaffected by cadence changes.  Marsh and Martin (1995) used an electronically braked bicycle ergometer to examine changes in EMG activity of five muscles of the lower limbs at five different cadences (50, 65, 80, 95, 110 rpm) at a constant power output of 200 Watts. Their findings agreed with those of Ericson et al. (1985): medial gastrocnemius EMG activity increased linearly with increases in cadence while soleus activity remained independent of cadence changes. Also, the peak amplitude of both soleus and medial gastrocnemius were found to occur earlier in the crank cycle with increasing cadence. This, Marsh and Martin explained as resulting from an increasing need to compensate for the electromechanical delay in the generation of muscle tension with increasing pedalling rate.  Sanderson et al. (2006) used a protocol similar to that of Marsh and Martin (1995) and, again found a systematic increase in the EMG activity of the medial gastrocnemius with increases in cadence, and no change in the EMG activity of the soleus. Despite being reported by several authors, the mechanism responsible for this phenomenon remains unclear. Several animal studies examining the electromyographic response of the muscles of the triceps surae during activities including walking, running, swimming, and paw shaking have found results that 27  agree with those from the studies in humans using a bicycle ergometer: the surface EMG activity from gastrocnemius generally increased with movement speed while that of soleus decreased or remained constant (Walmsley et al., 1978; Gardiner et al., 1982; Hodgson, 1983; Hutchison et al., 1989; Pierotti et al., 1989; Roy et al., 1991). The mechanism most frequently attributed to the differential effect of movement speed on EMG activity in these studies is differences in muscle fibre composition between the soleus and gastrocnemius.  Sanderson et al. (2006) also described evidence of load sharing between the two plantarflexor muscles that was dependent on pedaling rate. Both muscles exhibited a local peak in excitation that occurred in the second half of the pedal cycle (after the pedal had past bottom dead centre). As cadence increased, the second peak in medial gastrocnemius activity decreased, while the second peak in soleus activity increased, reciprocally. A post-hoc examination of the results from Marsh & Martin (1995) revealed a similar pattern of reciprocal changes in EMG activity.  2.0 Statement of the Problem  To the author’s knowledge, the only investigation into the effect of seat height on lower-limb EMG during cycling was a Master’s thesis by Amoroso (1994). Further research into the effects of lowering seat height would yield a greater understanding of muscle function, applicable not only to cycling, but to human movement in general. The first objective of this study was to expand upon the findings of Amoroso (1994) by examining the effect of decreased seat height on five additional muscles and at four additional cadences.  Previous research investigating the effect of cadence on muscle activity has been focused mainly on the muscles of the triceps surae. Insight into the cause of the changes in triceps surae EMG with changes in cadence, and as to why the effect is different in the soleus than in the gastrocnemius, might be gained by examining the effect of cadence on activation of additional muscles, including those of upper leg. The second objective of this study was to expand upon what is known of the effect of cadence on the activation of the muscles of the lower limb by examining the simultaneous response of eight upper- and lower-leg muscles, including the medial and lateral heads of the gastrocnemius, which have not been previously investigated together in the same study.  The effect of cadence on the activation of the triceps surae has been investigated in several studies; however, the factor or factors responsible for the effect of cadence on the activation of the triceps surae remain unclear. Several possible factors can be identified from what is already known of the effect of cadence, yet none of the factors have been formally investigated. One strategy that might help to uncover the role that these factors play is to 28  introduce a broad change in the kinematic requirements of cycling and to investigate whether changes in these factors correspond to changes in the effect of cadence on triceps surae EMG. Lowering the seat height has been shown to cause changes in many kinematic variables associated with cycling, including but not limited to joint angles, muscle lengths and muscle velocities (Too, 1990). The third objective of this study was to combine manipulations of seat height and cadence in an attempt to shed light on possible factors responsible for the effect of cadence on EMG from the muscles of the triceps surae.  3.0 Methods The participants of this study were graduate students from the University of British Columbia. All were physically active, had experience in cycling at least at the recreational level and were able to cycle for at least one hour at a moderate workload with minimal discomfort. Participants were made fully aware of the experimental protocol prior to participating in the study. All experimental procedures conformed to the ethical guidelines of the University of British Columbia and to the Declaration of Helsinki.  3.1 Data Collection  Eleven male participants rode a standard bicycle mounted on an electronically braked cycle ergometer (Schwinn Velodyne, Chicago, IL), which simulates the inertial characteristics of road riding and varies the resistive force at the rear wheel to maintain a constant power output with changes to cadence (Attaway et al., 1992). An optical encoder with 1024 steps was used to record crank angle and the top dead centre crank position (TDC). A cadence monitor (Cateye, USA) attached to the crank provided feedback to the participants who were instructed to keep their cadence as close as possible to one of the target cadences. The cranks were equipped with cleated pedals and toe clips. Participants wore appropriately fitted cycling shoes of identical make and model (Nike, model 8155C) that were specifically designed for the pedals used in this study. Heart rate was monitored but not collected with a telemetric heart rate monitor (Polar, USA) to ensure adequate recovery between trials. A 10-minute warm-up period was given during which participants cycled at 100 W with a cadence of 80 rpm. Following the warm-up period, participants cycled at each of two seat height conditions for a minimum of five minutes at each of five cadences (50, 65, 80, 95 and 110 rpm) at a constant nominal power output of 200 W. EMG data and sagittal-plane video were recorded during the final 30 seconds of each trial. Participants were not made aware of the exact timing of the 30-second data collection period. A minimum rest period of three minutes was 29  given between each trial, or until the heart rate returned to its approximate pre-trial value. A rest period of this length was chosen based on the finding that an increase in amplitude of EMG activity from the quadriceps muscle group due to localized fatigue after cycling at intensities as high as 100% VO2 max returned to pre-fatigue levels following three minutes of inactivity (Petrofsky, 1979). Since participants in the current study cycled at intensities much lower than 100% VO2 max, a three-minute rest period was assumed to be sufficient to allow full recovery between trials and to prevent muscular fatigue from having a confounding effect on EMG activity. The heart rate data supported this assumption. In one block of trials, the seat height was set to 100% of trochanteric height; this is the distance measured from the ground to the greater trochanter of the femur with the participant standing upright in bare feet with the knees fully extended and has been found to be the seat height associated with the lowest oxygen consumption (Shennum & deVries, 1976; Nordeen-Snyder, 1977). The seat height was defined as the linear distance from the centre of the pedal spindle to the highest point on the top surface of the seat along a straight line formed by the crank, seat tube and seat post (Hamley & Thomas, 1967). In the other block of trials, the seat height was set to 90% of trochanteric height. Each block of seat height trials, as well as the cadence manipulations within each block was completed in random order. The magnitude of the seat height manipulations were chosen according to the following rationale: setting the seat height to 100% trochanteric height provided a standardized control condition. The 90% condition was chosen to achieve the largest reduction in seat height while still allowing participants to complete the pedal cycle without both the knee and ankle joints reaching maximal flexion and dorsiflexion, respectively. With these seat height conditions, there were ten individual 5-minute trials for a total cycling time of 50 minutes.  Surface electromyographic (EMG) data were collected at a sampling frequency of 1200 Hz through a 12-bit analogue-to-digital converter using bipolar silver-silver chloride surface electrodes (1.5 cm center-to-center, Therapeutics Unlimited, USA) attached to the skin with double-sided adhesive pads over the following muscles of the left leg: soleus, lateral gastrocnemius, medial gastrocnemius, tibialis anterior, semimembranosus, vastus lateralis, rectus femoris and gluteus maximus. The electrodes were placed parallel to the presumed direction of the muscle fibres in the locations described by Basmajian and De Luca (1985). The electrodes provided a pre-amplification (gain = 35) at the recording site (CMRR = 87 dB at 60 Hz). Before placing the electrodes, the skin sites were shaved, abraded and cleaned with isopropyl alcohol to reduce source impedance. The electrodes, together with the pre- amplifier, were further secured to the skin with perforated tape to reduce motion artifacts. 30   Kinematics of the lower-limb segments were recorded simultaneously with the EMG data at 60 Hz using a video camera (Panasonic, WDV 5100, USA) positioned 6 m from the participant, with the lens axis oriented orthogonally to the rider’s sagittal plane. A reflective marker was placed over each of the anterior superior iliac spine (ASIS), posterior superior iliac spine (PSIS), greater trochanter (HIP), the estimated centre of rotation of the knee joint (KNEE), the lateral malleolus (ANKLE), the base of the calcaneus (HEEL) and the head of the fifth metatarsal (TOE) prior to data collection. Out-of-plane motion of the lower-limb was restricted by the riders’ cleated cycling shoes and toe clips.  3.2 Data Processing EMG data from the first twenty complete pedal cycles (recorded in the final 30 seconds of each 5-minute trial) were full-wave rectified and filtered (2nd order dual-pass Butterworth filter, 5 Hz cut-off frequency) to create a linear envelope. These data were separated into twenty individual pedal cycles using the location of the time- matched TDC pulse to determine the beginning and end of each cycle. The EMG data were then scaled in amplitude to the highest value of the linear envelope of the 50 rpm 100% seat height condition for each respective muscle. It is at this point that the first dependent measure was calculated: the integral with respect to time was determined for each of the twenty cycles before being averaged across cycles and participants. The amplitude-normalized EMG data were then time-normalized to 100 points. A post-hoc analysis revealed that participants were able to control their pedaling cadence to within ± 1 rpm (see section 4.0); therefore, the 100-point amplitude-normalized EMG array corresponded to 100% of the crank cycle in equal steps of 1%. This EMG array was averaged across the twenty pedal cycles for each participant and two additional dependent measures were calculated: the peak EMG value and the crank angle at which this value occurred. Video records for the same pedal cycles for which EMG data were recorded were digitized using the Peak Performance Technologies software package (Motus, version 7.1, USA). The two-dimensional coordinates for each reflective marker were calculated and filtered using a 2nd order, dual-pass Butterworth filter. The cutoff frequency that returned the highest signal-to-noise ratio for the movement of each marker along each axis was determined by examining the residuals using cut-off frequencies ranging from 1 to 10 Hz (Winter, 1990) (see table 3). These data were then time-matched to the corresponding EMG data file using a manually generated square-wave pulse, recorded as one channel of the A/D. This pulse also generated a white square on the video signal using an event 31  synchronization unit. The kinematic data arrays were made equal in length to the EMG arrays from one TDC to the next via linear interpolation. Table 3: Cut-off frequencies for kinematic data filtering  Cut-off Frequency (Hz) Marker X-axis Y-axis ASIS 3 3 PSIS 3 3 GT 3 3 KNEE 4 4 ANKLE 5 5 HEEL 5 5 TOE 5 5  Marker kinematics were used to compute segment lengths and joint angles of the hip, knee and ankle. The operational definitions of the segment lengths and joint angles are listed in table 4.  Table 4: Operational definitions of segment lengths and joint angles.  Segment Lengths  The rectilinear distance between the…  Thigh: HIP and the KNEE Shank: KNEE and the ANKLE Foot: ANKLE and the TOE  Joint Angles  The angle, measured counter-clockwise from the…  Hip: thigh segment to a vector directed from the hip joint centre to the ASIS Knee: thigh segment to the shank segment Ankle: foot segment to the shank segment  Muscle-tendon unit lengths over the pedal cycle were estimated using equations developed by Hawkins and Hull (1990) from averaged cadaveric muscle attachment data. These equations present muscle-tendon unit lengths as a percent of segment length. The velocities of muscle-tendon unit shortening and lengthening were calculated by differentiating the muscle length data using the finite difference central approximation. The sign of the muscle velocity data allowed periods of concentric and eccentric activity to be identified. Joint angle, muscle length, and muscle velocity data throughout the 30-second collection period were divided into twenty pedal cycles based on the 32  location of TDC in each of the data arrays. The data were time-normalized, and averaged across the twenty cycles. From these data, maximum, minimum, and ranges were computed for statistical analysis.  3.3 Statistical Analysis  A 5 (cadence) x 2 (seat height) repeated measures ANOVA was performed on each dependent variable. Mauchly’s test of sphericity was performed on each dependent measure to verify that the assumption of compound symmetry had been met. Greenhouse-Geisser epsilon values were used to adjust the F-value of dependent measures that did not meet the assumption of compound symmetry. The Greenhouse-Geisser method has been shown to over- estimate the degree to which the F-value is adjusted in cases where the epsilon value is greater than 0.7 (Stevens, 2002); therefore, when the Greenhouse-Geisser epsilon was found to be greater than 0.7, the Huynh-Feldt correction, a less conservative method, was used to adjust the F-value. For all statistical tests, a p-value of less than 0.05 was considered statistically significant. In variables yielding a cadence-by-seat height interaction, the influence of seat height on the relationship between cadence and the dependent measure in question was investigated by comparing the means within the cadence conditions using Repeated contrasts; this type of contrast test allows consecutive pairs of levels to be compared, for example level 1 with level 2, level 2 with level 3, level 3 with level 4, and so forth. Previous research has reported a linear relationship between cadence and many lower limb kinematic and electromyographic variables, making Repeated contrasts the best suited method of comparing means for these data (Sanderson et al., 2006). In variables yielding a main effect, data were pooled across the levels of other factor and Repeated contrasts were again used to further explore the main effect. A trend analysis was performed on each dependent measure, up to a maximum of a fourth-degree polynomial, to further describe its relationship with changes in cadence.  4.0 Results   During each trial, participants were asked to keep their cadence as close as possible to one of five target cadences. Table 5 presents the mean cadence for each trial, averaged across all subjects and both seat heights. These data indicate that participants were successful at matching the target cadence with little variability.   33  Table 5: Mean cadence for each trial, averaged across all participants and both seat height conditions  Target Cadence 50 rpm 65 rpm 80 rpm 95 rpm 110 rpm Actual Cadence (± st. dev.) 51.4 ± 1.3 65.9 ± 0.9 80.9 ± 0.9 95.7 ± 0.7 110.7 ± 0.9  4.1 Joint Angle 4.1.1 Ankle Angle  Figures 1A and 1B present the pattern of ankle motion for both seat-height conditions. A lowered seat height resulted in a compression of the range of ankle motion. A significant seat height-by-cadence interaction was found for the maximum ankle angle and the range of ankle motion (Max: F(4,40)=4.184, p<0.001; Range: F(4,40)=3.019, p=0.029). Repeated contrasts revealed that, in the high-seat condition, the range of ankle motion decreased linearly with increasing cadence until the 95 rpm condition, beyond which no further decrease in range of motion was seen. In the low-seat condition, ankle range of motion decreased linearly until the 80 rpm condition, beyond which it increased with further increases in cadence. The changes in the range of motion were driven by the maximum degree of plantarflexion. 4.1.2 Knee Angle  Changing the seat height resulted in a similar compression of the knee joint angle profile, as shown in Figures 2A and 2B. Main effects of seat were found for the minimum, maximum and range of knee angles, indicating that the knee operated over a smaller range of motion and from a position of increased flexion in the low- seat condition (Min: F(1,10)=555.482, p<0.001; Max: F(1,10)=328.950, p<0.001; Range: F(1,10)=161.360, p<0.001). A main effect of cadence was also found for the knee angle minimum, maximum and range (Min: F(4,40)=21.051, p<0.001; Max: F(4,40)=29.682, p<0.001; Range: F(4,40)=18.394, p<0.001). Repeated contrasts revealed that as cadence increased, the knee operated over a smaller range of motion and from a position of increased flexion; these changes occurred with a linear trend (Linear Trends - Min: F(1,10)=29.536, p<0.001; Max: F(1,10)=50.121, p<0.001; Range: F(1,10)=36.448, p<0.001). 4.1.3 Hip Angle   The range of hip joint motion was less at the low-seat height, as indicated by a significant main effect of seat  (F(1,10)=101.156, p<0.001) (see Figures 3A and 3B). The decrease in range of motion was a result of a decrease in the maximum hip angle (more flexed) and an increase in the minimum hip angle (less flexed). A main effect of 34  cadence for hip joint range of motion was also found (F(4,40)=9.199, p<0.001). The range of motion was greatest at the 80 rpm condition and was less at cadences greater and less than 80 rpm; these changes followed a quadratic trend (Quadratic Trend - F(1,10)=47.759, p<0.001). 4.2 Muscle Length 4.2.1 Soleus Length  Given that the soleus muscle is mono-articular, crossing only the ankle joint, changes in soleus length across the crank cycle were directly reflective of changes in ankle angle. Thus, the statistical results obtained for the kinematics of the ankle angle would be expected to be similar to those for the soleus length.  Figures 4A and B present the soleus length across the crank cycle for each cadence and seat-height condition. Soleus length changed differentially with increases in cadence between the two seat-height conditions, as indicated by a significant seat height-by-cadence interaction (Min: F(4,40)=4.56, p=0.004; Range: F(4,40)=3.449, p=0.016). Repeated contrasts revealed that, in the high-seat condition, the range of soleus motion decreased linearly with increasing cadence until the 95 rpm condition, beyond which no further decrease in range of motion was seen. In the low-seat condition, soleus range of motion decreased linearly until the 80 rpm condition, beyond which it increased with further increases in cadence. The changes in the range of motion were driven by the minimum soleus length. 4.2.2 Medial Gastrocnemius Length Figures 5A and B present the medial gastrocnemius length across the crank cycle for each cadence and seat- height condition. No statistical difference was found in the minimum and maximum lengths of medial gastrocnemius between the two seat heights (Min: F(1,10)=0.001, p=0.971; Max: F(1,10)=2.084, p=0.179). As cadence increased, the medial gastrocnemius operated at a shorter absolute length and over a smaller range of motion. This was driven by a decrease in the maximum length, as indicated by a main effect of cadence (Max: F(4,40)=20.896, p<0.001; Range: F(4,40)=10.370, p=0.001). Repeated contrasts revealed significant differences between each consecutive pair of cadences with the exception of the 80 and 95 rpm conditions. The changes in maximum length and range of motion followed a linear trend (Max: F(1,10)=40.443, p<0.001; Range: F(1,10)=15.572, p=0.003).   35  4.2.3 Lateral Gastrocnemius Length  Figures 6A and B present the lateral gastrocnemius length across the crank cycle for each cadence and seat- height condition. The medial and lateral gastrocnemii have proximal insertions that are similar in rectilinear distance from their common distal insertion at the Achilles tendon. Thus, little difference was expected in the changes in the muscle’s lengths. No statistical difference was found in the minimum and maximum lengths of lateral gastrocnemius between the two seat heights (Min: F(1,10)=2.696, p=0.132; Max: F(1,10)=3.139, p=0.107). As cadence increased, the lateral gastrocnemius operated at a shorter absolute length and over a smaller range of motion. This was driven by a decrease in the maximum length, as indicated by a main effect of cadence (Range: F(4,40)=11.270, p=0.002; Max: F(4,40)=21.045, p<0.001). Repeated contrasts revealed significant differences between each consecutive pair of cadences with the exception of the 80 and 95 rpm conditions. The changes in maximum length and range of motion followed a linear trend (Max: F(1,10)=40.170, p<0.001; Range: F(1,10)=14.946, p=0.003). 4.2.4 Tibialis Anterior Length  The tibialis anterior is a mono-articular antagonist to the triceps surae. As such, the statistical results obtained for the tibialis anterior length were expected to be similar in magnitude and opposite in direction to those for the soleus length.  Tibialis anterior length changed differentially with increases in cadence between the two seat-height conditions, as indicated by a significant seat height-by-cadence interaction (Max: F(4,40)=4.57, p=0.004; Range: F(4,40)=3.449, p=0.016) (see Figures 7A & B).  Repeated contrasts revealed that, in the high-seat condition, the range of tibialis anterior motion decreased linearly with increasing cadence until the 95 rpm condition, beyond which no further decrease in range of motion was seen. In the low-seat condition, tibialis anterior range of motion decreased linearly until the 80 rpm condition, beyond which it increased with further increases in cadence. The changes in the range of motion were driven by the maximum tibialis anterior length. 4.2.5 Semimembranosus Length  Figures 8A and B present semimembranosus length across the crank cycle for each cadence and seat-height condition. Semimembranosus operated at a shorter length and over a smaller range of motion in the low-seat condition (Range: F(1,10)=125.953, p<0.001). This was due to a decrease in both the maximum and minimum lengths (Max: F(1,10)=240.406, p<0.001; Min: F(1,10)=119.927, p<0.001). As cadence increased, semimembranosus operated over a smaller range of lengths due to a linear decrease in the maximum length (Max: F(4,40)= 14.456, p<0.001; 36  Range: F(4,40)= 22.639, p<0.001; Linear Trend: F(4,40)= 26.863, p<0.001). Repeated contrasts revealed significant differences between 65 and 80 rpm conditions and between the 95 and 110 rpm conditions. The effect of cadence on semimembranosus length was independent of changes in seat height. 4.2.6 Vastus Lateralis Length  Vastus lateralis operated at a longer length and over a smaller range of motion in the low-seat condition, independent of cadence (Range: F(1,10)=145.269, p<0.001) (see Figures 9A & B). This was due primarily to an increase in the minimum length in the low-seat condition (Max: F(1,10)=9.183, p=0.013; Min: F(1,10)=145.117, p<0.001). The minimum length occurred at the 50 rpm cadence at both seat heights. As cadence increased, the minimum length increased linearly while the maximum length remained constant; therefore, as cadence increased, the range of lengths decreased. Repeated contrasts  performed on the range of lengths revealed significant differences between each consecutive pair of cadences conditions. These changes with increasing cadence were independent of changes in seat height (Min: F(4,40)= 28.883, p<0.001; Range: F(4,40)= 28.944, p<0.001). 4.2.7 Rectus Femoris Length  The length of the rectus femoris is dependent on changes in the hip and knee angles. Despite being a bi- articular muscle, the overall shape of its length-vs-crank angle profile changed little with changes in seat height; however, it operated at a greater length  and over a smaller range of lengths in the low-seat condition (Min: F(1,10)=299.857, p<0.001; Max: F(1,10)=151.7, p<0.001; Range: F(1,10)=136.289, p<0.001) (see Figures 10A & B).  At both seat heights, rectus femoris operated at the shortest length and over the greatest range of lengths at the 50 rpm cadence. As cadence increased, the muscle operated at a greater length and over a smaller range of lengths (Min: F(4,40)=17.082, p<0.001; Max: F(4,40)=3.753, p=0.046; Range: F(4,40)=21.372, p<0.001). These changes in rectus femoris length yielded a linear trend (Range: F(1,10)=52.217, p<0.001). Repeated contrasts performed on the range of lengths revealed significant differences between the 50 and 65 rpm conditions and between the 65 and 80 rpm conditions. 4.2.8 Gluteus Maximus Length  Gluteus maximus length was not statistically different between the two seat heights (Min: F(1,10)=0.710, p=0.419; Max: F(1,10)=0.264, p=0.619), nor between the five cadences (Min: F(4,40)=1.505, p=0.249; Max: F(4,40)=0.999, p=0.371) (see Figures 11A & B). 37  4.3 Muscle Velocity The shortening and lengthening velocities for all muscles increased with increasing cadence, with the exception of the medial and lateral gastrocnemii (see Figures 4-11C & D). At each cadence, the shortening and lengthening velocities were less at the low seat height. 4.4 EMG The calculation of muscle velocity allowed the EMG data to be parsed into those associated with concentric muscle actions and those associated with eccentric muscle actions. Changes in the direction of muscle action and the magnitude of muscle activation will be discussed in this section. The statistical results from the EMG data for all muscles are summarized in Table 6. Table 6: Summary of statistical results of a 2 (seat) by 5 (cadence) repeated measures ANOVA for the peak and integral EMG values for each muscle. 3 indicates a main effect; – indicates an interaction; 2 indicates non-signifcance.   Peak Integral  Interaction Seat Cadence Interaction Seat Cadence Soleus 2 2 2 2 2 2 Medial Gastrocnemius 2 3 3 2 3 3 Lateral Gastrocnemius 2 2 3 3 – – Tibialis Anterior 2 2 2 3 – – Semimembranosus 2 2 3 2 2 3 Vastus Lateralis 2 2 3 2 2 3 Rectus Femoris 3 – – 2 2 2 Gluteus Maximus 2 3 3 2 2 3  4.4.1 Soleus EMG  The linear envelope of normalized soleus EMG across the crank cycle displayed one peak at a crank angle of, on average, 90° (see Figures 4E & F and 12A-E). Minimal activity was seen during the recovery phase (from 180° to 360°).  Soleus activation was insensitive to changes in cadence and seat height. An interaction between seat height and cadence was not found for the soleus EMG integral (F(4,40)=0.468; p=0.585), nor for the peak EMG (F(4,40)=1.020; p=0.369). Main effects of cadence and seat height were not found for both the integral (cadence: F(4,40)=2.944, p=0.092; seat height: F(1,10)=0.293, p=0.600), and peak EMG values (cadence: F(4,40)=1.513, p=0.245; seat height: F(1,10)=1.682, p=0.224). 38   The timing of peak soleus EMG, expressed as a position along the crank cycle, yielded a seat-by-cadence interaction (F(4,40)=6.577, p<0.001). Repeated contrasts revealed that, in the high-seat condition, the timing of peak EMG occurred earlier in the crank cycle with increases in cadence, whereas, in the low-seat condition, the timing of peak EMG did not change with increases in cadence. 4.4.2 Medial Gastrocnemius EMG  The linear envelope of normalized medial gastrocnemius EMG displayed one peak that occurred at a crank angle between 90-180°, greater in magnitude than the activation over the rest of the crank cycle. This peak occurred slightly later in the crank cycle than the largest peak in soleus EMG (see Figures 5E & F and 13A-E). This was consistent at all cadences and both seat heights.  Main effects of seat height were found for the peak and integral values (Peak: F(1,10)=92.871, p<0.001; Integral: F(1,10)=121.188, p<0.001). When these data were averaged across all cadences, the EMG integral from the low-seat condition was 65% of that from the high-seat condition. Peak EMG in the low-seat condition was an average of 62 % of that in the high-seat condition.  Main effects of cadence were found for the peak and integral values (Peak: F(4,40)=29.931, p<0.001; Integral: F(4,40)=22.321, p<0.001). Repeated contrasts revealed significant differences between each consecutive pair of means. A trend analysis revealed that peak EMG increased linearly with increasing cadence (Linear Trend: F(1,10)=32.409, p<0.001). The EMG integral also increased with increasing cadence, but did so only at cadences greater than 65 rpm.  The timing of peak EMG in medial gastrocnemius yielded main effects of both seat height (F(1,10)=21.272, p=0.001) and cadence (F(4,40)=16.616, p<0.001). When averaged across all cadences, peak EMG occurred 35° earlier in the crank cycle in the high-seat condition than in the low-seat condition. When averaged across both seat heights, peak EMG occurred latest in the 50 rpm condition. As cadence increased, peak EMG occurred earlier in the crank cycle, with its earliest position being seen in the 110 rpm condition. 4.4.3 Lateral Gastrocnemius EMG  Lateral gastrocnemius displayed two peaks that were similar in amplitude, the first peak occurring in the second quarter of the crank cycle and the second peak occurring in the third quarter of the crank cycle, well after the crank had passed the bottom dead centre position (see Figures 6E & F and 14A-E). In the high-seat condition, the two peaks remained intact with increases in cadence, despite overall increases in the EMG amplitude. In the low- seat condition, the two peaks were most distinct at the highest cadence and moved centrally as cadence decreased 39  until, at the 50 rpm cadence, there was only a single peak of EMG activity. The two merging peaks did not sum to yield a single peak of greater magnitude: the maximum value of the single peak was not different from the maximum value of the double peaks.  A significant interaction was found for the EMG integral (F(4,40)=6.243, p=0.001). Repeated contrasts revealed that in the high-seat condition, only the 110 rpm cadence was significantly greater than the previous cadence, whereas in the low-seat condition, cadences of 80 rpm and higher were greater than the previous cadence. The EMG integral values from the lateral gastrocnemius at each cadence and seat-height condition, presented in Figure 15, suggest that a low seat height decreased lateral gastrocnemius activity, but that at high cadences, the effect of the low seat was overcome by the sensitivity of lateral gastrocnemius to increases in cadence. The peak EMG from lateral gastrocnemius did not yield a significant seat-by-cadence interaction (F(4,40)=2.257, p=0.08).  A main effect of seat height was not found for peak EMG of the lateral gastrocnemius (F(4,40)=0.382, p=0.079). The amplitude of the first peak decreased while that of the second peak increased in the low-seat condition compared to the high-seat condition, yet the global maximum remained unchanged. The double-peaked landscape of the lateral gastrocnemius normalized EMG linear envelope was not as sensitive to changes in seat height as the single-peaked landscape of medial gastrocnemius.  The timing of two peaks of the lateral gastrocnemius EMG linear envelope did not change with changes in cadence (F(4,40)=1.288, p=0.298). In the low-seat condition, the central migration of the two peaks with decreasing cadence caused apparent changes in the timing of peak EMG. 4.4.4 Tibialis Anterior EMG  The linear envelope of normalized tibialis anterior EMG across the crank cycle displayed a large peak between a crank angle of 270° and 360° (see Figures 7E & F and 16A-E). A second, smaller peak was also seen at approximately 180° when the crank was close to the bottom dead centre position.  A significant seat-by-cadence interaction was found for the tibialis anterior EMG integral (F(4,40)=6.288, p=0.007). In the high-seat condition, the integral value increased linearly with increases in cadence (Linear Trend: F(1,10)=7.167, p=0.023). In the low-seat condition, the integral value at the 50 rpm cadence was greater than that at the 65 rpm cadence and did not begin to increase with increasing cadence until the 80 rpm condition. The integral values, presented in Figure 17, suggest that a decrease in seat height decreased the effect of cadence on tibialis anterior EMG, specifically at low cadences. 40   The timing of the large peak in tibialis anterior EMG did not change with increasing cadence (F(4,40)=0.472, p=0.667), nor with a decrease in seat height (F(1,10)=4.746, p=0.054). The small peak occurred earlier with increasing cadence (F(4,40)= 4.431, p=0.031), and in the high-seat condition (F(1,10)= 4.055, p=0.028). 4.4.5 Semimembranosus EMG  The linear envelope of normalized semimembranosus EMG across the crank cycle revealed one peak between 90° and 180° (see Figures 8E & F and 18A-E). A main effect of cadence was found for both the peak and integral values. Using Repeated contrasts for both the peak and integral values, significant differences were found only between the 95 and 110 rpm conditions. A trend analysis revealed that changes in peak and integral values were linear (Peak: F(4,40)= 16.783, p<0.001; Integral: F(4,40)= 8.988, p=0.001). There was no statistical difference in the peak or integral values of semimembranosus EMG between the two seat heights (Peak: F(1,10)= 2.887, p=0.124; Integral: F(1,10)= 2.132, p=0.175). As cadence increased, peak semimembranosus EMG occurred slightly earlier in the crank cycle (F(4,40)= 4.094, p= 0.045). 4.4.6 Vastus Lateralis EMG  The linear envelope of normalized vastus lateralis EMG across the crank cycle revealed one peak between 0° and 90° (see Figures 9E & F and 19A-E). A main effect of cadence was found for both the peak and integral EMG values (Peak: F(4,40)= 3.621, p= 0.042; Integral: F(4,40)=2.956, p= 0.031). Repeated contrasts revealed that, for the integral value, a significant difference was found only between the 95 and 110 rpm conditions, whereas for the peak value, a significant difference was found only between the 80 and 95 rpm conditions. The changes in both the peak and integral values followed a quadratic trend (Peak: F(4,40)= 8.168, p<0.019; Integral: F(4,40)= 6.099, p=0.036). There was no statistical difference in the peak or integral values of vastus lateralis EMG between the two seat heights (Peak: F(1,10)= 1.866, p=0.205; Integral: F(1,10)= 1.977, p=0.188). 4.4.7 Rectus Femoris EMG  The linear envelope of normalized rectus femoris EMG across the crank cycle displayed two peaks: the first occurred prior to TDC, between 270° and 360°; the second occurred after TDC, between 0° and 90° (see Figures 10E & F and 20A-E). As cadence increased, the first peak decreased in amplitude, while the second peak increased in amplitude. This occurred at both seat heights. The EMG integral did not yield a main effect of cadence. The two peaks of rectus femoris activity occurred earliest in the crank cycle at the 110 rpm condition and occurred later with 41  decreasing cadence. There was no statistical difference in the peak or integral values of rectus femoris EMG between the two seat heights (Peak: F(1,10)= 0.391, p=0.547; Integral: F(1,10)= 1.473, p=0.253). 4.4.8 Gluteus Maximus EMG  The linear envelope of normalized gluteus maximus EMG across the crank cycle displayed one peak between 45° and 135° (see Figures 11A & B and 21A-E). Little activity was seen between the BDC and TDC positions.  Peak gluteus maximus EMG was lower in the low-seat condition, as indicated by a significant main effect of seat height (F(1,10)=5.614, p=0.039). There was no statistical difference in the integral value between the two seat heights (F(1,10)=0.502, p=0.495).  EMG activity from gluteus maximus was greatest at the 50 rpm cadence and decreased linearly with increasing cadence, as indicated by a main effect of cadence for the peak and integral EMG values (Peak: F(4,40)=3.537, p=0.015; Integral: F(4,40)=13.042, p=0.001). Repeated contrasts revealed significant differences between the 50 and 65 rpm conditions and between the 65 and 80 rpm conditions. Gluteus maximus was the only muscle to exhibit an effect of cadence in this direction. As cadence increase, peak gluteus maximus EMG occurred earlier in the crank cycle (F(4,40)=32.857, p<0.001). 5.0 Discussion  The three objectives of this study were described in section 2.0. Here, the three objectives are revisited and discussed in light of the results obtained. In section 5.1, the electromyographic response of a decrease in seat height is summarized and several possible physiological factor(s) responsible for the effect are discussed. In section 5.2, the electromyographic response to changes in cadence is summarized, and four factors that could potentially be responsible for such an effect are described. In section 5.3, the effects of changes in cadence on several kinematic variables are compared between the two seat heights, and evidence for or against each of the factors listed in section 5.2 is discussed. Sections 5.4-5.7 relate to topics not central to the three primary purposes of the study, yet merit formal discourse. 5.1 Objective I: Effect of Seat Height  The response of the soleus, medial gastrocnemius and tibialis anterior to a decrease in seat height was consistent with that reported by Amoroso (1994): soleus activity did not change, medial gastrocnemius activity 42  decreased and tibialis anterior activity increased. A surprising finding of the current study was that the lateral gastrocnemius, which was not investigated by Amoroso (1994), did not respond to a decrease in seat height in the same manner as the medial gastrocnemius. This was one of several differences found in the functioning of the medial and lateral heads of the gastrocnemius muscle. Given the similarity of the attachment sites of the medial and lateral gastrocnemius, kinematic changes in cycling were expected to exhibit similar effects on the two muscles. The differences between the gastrocnemii are discussed in greater detail in section 5.5.  Of the four additional muscles that were investigated (vastus lateralis, rectus femoris, semimembranosus and gluteus maximus), all but gluteus maximus, whose activation was less at the low seat height, failed to yield a main effect of seat height. Given that all experimental conditions were performed at a constant power output, it seems unlikely that gluteus maximus, one of the primary muscles involved in hip extension, would undergo a decrease in activation without a compensatory increase in activation from a knee extensor, vastus lateralis. To explore further the hypothesis of load sharing between the gluteus maximus and vastus lateralis, the difference in the EMG integral between the two seat heights was calculated at each cadence and compared between the two muscles (see Figure 22). A statistically significant difference between the two muscles was not found, although a trend towards a differential response between the two muscles with a decrease in seat height was seen.  The lack of change in activation of the vastus lateralis, rectus femoris and semimembranosus might suggest that the manipulation of seat height was simply not large enough to induce observable changes in the activation of these muscles; however, participants typically reported feeling as though their gluteus maximus and knee extensors were more active in the low-seat condition. It is possible that, with the seat height decreased, individual participants adopted different strategies of muscle activation to maintain a constant power output and that, when combined, the data did not yield a statistically significant result in any one direction.  Myoelectrical signals from the muscles of the triceps surae have been found to adapt to maximize muscle tension for a given muscle length (see section 1.1.2 for a review). Several authors have demonstrated, using both indwelling and surface EMG electrodes, that activation of the medial and lateral gastrocnemii decreases when the knee is in a flexed position and have attributed this to a decrease in muscle length (Cresswell et al., 1995; Tamaki et al., 1997; Miaki et al., 1999; Kennedy & Cresswell, 2001; Shinohara et al., 2006). In the current study, the decrease in medial gastrocnemius EMG in the low-seat condition appears, at first glance, to be in agreement with the findings of these studies: in the low-seat condition, the knee was in a position of greater flexion which, if the ankle angle were held constant, would cause the medial gastrocnemius to shorten. However, with the seat height decreased, the 43  ankle was more dorsiflexed, increasing the length of the medial gastrocnemius, resulting in no net change in muscle length between the two seat heights.  If it was not the length of the medial gastrocnemius that stimulated the decrease in its activation, what was it? In each of the studies mentioned previously, length of the gastrocnemii was manipulated by changing the knee angle while the ankle angle was held constant. Any change in gastrocnemius activation that was associated with decreased muscle length would also have been associated with increased knee flexion. This might suggest that it was knee angle, not muscle length that was responsible for the decrease in gastrocnemius EMG. Before expanding these findings to the current study, it is important to note the minimum change in knee angle required to illicit a decrease in medial gastrocnemius activation. Beginning from a position of full knee extension, the smallest magnitude of knee flexion found to have led to a decrease in medial gastrocnemius EMG was 90° which is greater than the difference in the mean knee angle seen between the two seat heights in the current study ( 20°). Cresswell et al. (1995) was the only study to examine the effect of a change in knee angle of less than 90°; they reported finding a trend of decreasing medial gastrocnemius EMG with decreases in knee angle as small as 30°, although this change in EMG did not reach statistical significance until a difference in knee angle of 90°. Despite increases in knee flexion that were less than the minimum previously associated with changes in EMG, a greater than 60% reduction in gastrocnemius EMG was seen in the current study. This was likely due to the greater number of muscles involved in pedalling a cycle ergometer versus the number involved in isometric plantarflexion. With more muscles acting as synergists during cycling, there is greater opportunity for load sharing, and gastrocnemius activity can be reduced to a greater extent without a decrease in torque output.  Inhibitory pathways from joint receptors in the knee seem the most plausible physiological structure involved in reducing gastrocnemius activity with changes in knee angle. Feedback from knee joint receptors has been found to have an inhibitory influence on the quadriceps muscle group (Stokes & Young, 1984; Young et al., 1987). A second possibility for the stimulus location, put forth by Tamaki et al. (1997), is via heteronymous connections from muscle spindles in the ipsilateral mono-articular knee extensors (ie: vastus medialis, vastus lateralis and vastus intermedius). Signals from these two groups of proprioceptors (knee joint receptors and heteronymous muscle spindles) would provide information regarding changes in gastrocnemius length that were due to increased knee flexion. Although information regarding both the knee and ankle angles is required to determine the net length of the gastrocnemius, it is likely that the knee angle is the more valuable of the two joint angles when coordinating load sharing among the triceps surae, since any change in length of the gastrocnemius due to changes 44  in ankle angle will also be experienced by the soleus. In other words, only changes in knee angle will differentially affect the lengths of the gastrocnemius and soleus and, therefore, knee angle alone may be the stimulus driving the derecruitment of the gastrocnemius. Results reported by Arampatzis et al. (2006) support the hypothesis that knee angle is responsible for the decrease in gastrocnemius EMG. In their study, medial gastrocnemius fascicle length and surface EMG were measured during isometric maximal voluntary plantar flexion contractions at various combinations of ankle and knee angles. It was found that surface EMG decreased with increasing knee flexion despite there being no change in fascicle length.  An assumption of the knee angle hypothesis is that, for a given change in ankle angle, the soleus and gastrocnemius will experience an equal change in their tension-producing capabilities. For example, if, while the knee angle was held constant and the ankle was placed in a position of plantarflexion, the gastrocnemius shifted to a non-optimal length, information regarding the ankle angle would prove important in mediating the derecruitment of the gastrocnemius, and knee angle alone would not provide enough information. Maganaris et al. (2001; 2003), mapped the length-tension curves of the in vivo human soleus and gastrocnemius and found that, with the knee fully extended and the ankle at angles ranging from 30° of dorsiflexion to 30° of plantarflexion, the two muscles operated along the ascending limb. Only at 45° of plantarflexion did the soleus reach the plateau portion of the curve. The magnitude of ankle angles reached in the current study ranged from 22° of plantarflexion to 39° of dorsiflexion. Within this range, the gastrocnemius and soleus would have operated almost entirely along the ascending limb. The difference in the mean ankle angle between the two seat heights in the current study was only 6°. This suggests that the reduction in tension-producing capabilities with changes in ankle angle, alone, was similar between the two muscles. Further information regarding the slope of the ascending limb of each muscle would provide a better estimate of the difference.  As was the case for the medial gastrocnemius, gluteus maximus length did not differ between seat heights. The decrease in gluteus maximus activity with the decrease in seat height could, therefore, not be attributed to a change in muscle length. The velocity profile of the gluteus maximus was also very similar between the two seat heights, implying that the change in EMG was not due to a shift along the muscle’s force-velocity curve.  As the length of a muscle’s moment arm decreases, the metabolic cost required to maintain a constant magnitude of muscle force increases (Lou et al., 1997; Ferguson et al., 2001). In the same way that the economy of muscle force production is maximized through the selective activation of synergists with optimal length-tension properties, a muscle’s moment-arm length may also be a factor mediating selective recruitment within a group of 45  synergists. The hip angle and angular velocity profiles across the crank cycle were similar in shape between the two seat heights. Since the gluteus maximus is a mono-articular muscle, crossing only the hip joint, it can be inferred from the similarities of the hip joint kinematics that the length of the gluteus maximus moment arm at the femur was also similar between the two seat height conditions. This suggests that a difference in moment arm length was not responsible for the decrease in gluteus maximus activation. It was likely not a single factor that led to the derecruitment of the gluteus maximus in the low seat height condition, but rather the combined effect of numerous influences, including properties of other lower limb muscles. Since the demand for cycling power output was constant, increases in the contribution of other muscles, which may not have necessarily been expressed as changes in EMG, may have allowed a decrease in the recruitment of the gluteus maximus.  The decrease in seat height led to a decrease in length and an increase in activation of the tibialis anterior, although this effect was somewhat hidden by the increase in EMG with increasing cadence in the high-seat condition. A decrease in tibialis anterior activation with an increase in ankle dorsiflexion has been shown previously by Vander Linden et al. (1991) and Pasquet et al. (2005). Pasquet et al. (2005) demonstrated that a change in ankle angle from 10° of plantarflexion to 10° of dorsiflexion caused the tibialis anterior to operate along the ascending limb of the length-tension curve rather than along the plateau. With this change in ankle angle, it was also found that the pennation angle of the tibialis anterior increased by 2.1° and 3.9° from the values measured at rest and during a maximal voluntary contraction, respectively (Pasquet et al., 2005).  Muscle length is only one of several factors, including moment-arm length and pennation angle that can limit the mechanical potential of a muscle to produce tension with changes in joint angle. Within a system of synergistic muscles, variations in muscle architecture will cause each of the mechanical factors to respond differently to changes in limb kinematics. Each variable must be taken into account when optimizing the economy of muscle function. There does not exist a single group of proprioceptors, muscle spindles for example, that can provide information on the state of each of these variables; therefore, it seems unlikely that changes in muscle excitation with changes in joint kinematics would be the result of feedback from a single afferent source. Instead, homonymous feedback from structures such as spindles is likely combined with both homonymous and heteronymous feedback from joint and cutaneous receptors and Golgi tendon organs to determine the overall force- generating potential of a muscle for a given kinematic state. 46  5.2 Objective II: Effect of Cadence  The effects of increasing cadence on activation of the soleus and medial gastrocnemius were consistent with those reported by Sanderson et al. (2006): soleus activity was insensitive to changes in cadence, while medial gastrocnemius activity increased systematically with increasing cadence. The data from the current study also agreed with those reported by Marsh & Martin (1995) for the soleus and gastrocnemii, as well as for the tibialis anterior, vastus lateralis and rectus femoris, which were not investigated by Sanderson et al. (2006). A novel finding of this study was that the medial and lateral gastrocnemii responded differently to increases in cadence.  The objective, in combining manipulations of seat height and cadence, was to provide some insight into the factor(s) responsible for effect of cadence on activation of the triceps surae. Analysis of the kinematic data yielded new information regarding the role of four factors: muscle length, muscle velocity, joint angle and the direction of muscle action. Before discussing evidence for or against these factors, it is important to understand how they could potentially influence muscle activation.  The degree of excitation required to produce a given amount of muscle tension changes with the muscle’s position along its length-tension and force-velocity curves (see sections 1.1 and 1.2). If the point at which the soleus and gastrocnemii operated along their respective length-tension and force-velocity curves were to change differentially with an increase in cadence, a differential change in excitation between the two muscles might be required to maintain the same degree of power output.  The changes in muscle activation that occurred with a decrease in seat height revealed that joint angle might play a more important role than muscle length in mediating the coordination of load sharing between synergistic muscles (see section 5.1).  Changes in joint kinematics with changes in cadence might also contribute to changes in muscle activation.  Many authors have reported differences in surface EMG between concentric and eccentric muscle actions of equal muscle tension and magnitude of velocity (see section 1.3.2). Relating specifically to the triceps surae, surface EMG activity from the soleus has been found to decrease with submaximal eccentric actions while that from gastrocnemius has been found to increase (Nardone & Schieppati, 1988). Differential changes in the proportion of concentric and eccentric muscle actions between the soleus and gastrocnemii with increases in cadence might contribute to the differential effect of cadence between these two muscles.  47  5.3 Objective III: Combined Manipulations  If a given factor were responsible for changing the degree of muscle activation with increases in cadence, it would be expected to do so in both seat height conditions. Therefore, evidence for or against the role of each factor can be determined by examining the consistency of its relationship with cadence between the two seat heights. Following the same reasoning, the role of a given factor in mediating the differential effect of cadence on the soleus and gastrocnemius can be assessed by examining the consistency of the factor’s relationship with changes in cadence between the two muscles; however, it is possible that different factors are responsible for the electromyographic response of each muscle. In this section, qualitative evidence for or against several factors will be discussed. 5.3.1 Muscle Length  A comparison of the changes in muscle length with increases in cadence at each seat height suggests that soleus length was not responsible for mediating the effect of cadence on soleus EMG. The relationship between soleus length and cadence was different between the two seat heights, whereas soleus EMG did not respond to changes in cadence at either seat height.  The role of gastrocnemius length in mediating the response of gastrocnemius activation to changes in cadence could not be refuted by the observed changes in gastrocnemius length and cadence at each of the seat heights. Medial and lateral gastrocnemius length and EMG decreased with increases in cadence, and these relationships were consistent between the two seat heights.  Comparing the relationship of muscle length and EMG with increasing cadence between the soleus and gastrocnemius provided evidence against the length of these muscles as factors responsible for the differential effect of cadence on the soleus and gastrocnemius EMG. The length of both the soleus and gastrocnemius decreased with increasing cadence, while the changes in EMG with increasing cadence were different between the two muscles. If the differential effect of cadence on soleus and gastrocnemius EMG were due to changes in muscle length, the response of muscle length to increases in cadence would be expected to be similar to the response of EMG for both muscles.  In section 5.1, the possibility was discussed that the stimulus mediating the effect of seat height on gastrocnemius EMG was located in the mono-articular knee extensor muscles. To investigate the possibility that the same factor was responsible for the effect of cadence, changes in the length of the vastus lateralis with increases in 48  cadence were compared to changes in gastrocnemius EMG with increases in cadence at each seat height. Vastus lateralis length was found to increase with increasing cadence, a trend that was seen in both seat height conditions, despite an overall reduction in vastus lateralis length at the low seat height (see Figures 9A & B). These changes in muscle length with changes in cadence and seat height were consistent with the changes seen in gastrocnemius EMG. As mentioned in section 5.1, the common insertion of the triceps surae via the Achilles tendon presented a situation in which information regarding changes in the length of the soleus and gastrocnemius, individually, may be of less importance in optimizing the economy of the muscle force generation than information regarding the ratio of gastrocnemius length-to-soleus length. The changes seen in vastus lateralis length support this notion, in that, as vastus lateralis length increased, the ratio of gastrocnemius length-to-soleus length decreased, and a reduction in gastrocnemius activation would be expected. The results of the current study suggest that vastus lateralis length cannot be refuted as a possible factor responsible for the changes in gastrocnemius EMG with changes in cadence and seat height and is a topic that merits further research. 5.3.2 Muscle Velocity  The velocity of soleus shortening and lengthening increased with increases in cadence to a greater extent in the high-seat condition than in the low-seat condition. This was not consistent with the relationship between soleus EMG and cadence. The effects of cadence on muscle velocity and EMG between the two seat heights were also inconsistent for both the medial and lateral gastrocnemius. Figure 23 presents the relationship between the maximal velocities of shortening with increasing cadence for the muscles of the triceps surae at each seat height. These data support the null hypothesis that changes in muscle activation with increases in cadence were not due to changes in the velocity of muscle action and could not explain the differential effect of cadence on activation of the soleus and gastrocnemius. 5.3.3 Joint Angle  With increases in cadence, the knee became increasingly flexed and operated over a smaller range of motion. These changes were consistent between the two seat heights, suggesting that knee angle cannot be refuted as a possible factor responsible for the changes in gastrocnemius EMG with increases in cadence.  Unlike the knee angle, the ankle angle responded differently to increases in cadence between the two seat heights which suggests that ankle angle was not a factor mediating the EMG response of the triceps surae to increases in cadence. 49  5.3.4 Direction of Muscle Action  To investigate the role of the direction of muscle action on the response of triceps surae activity to increases in cadence, the EMG integral across the pedal cycle was divided into that associated with muscle shortening and that associated with muscle lengthening. Changes in the ratio of concentric to eccentric EMG integral with changes in cadence were compared between the two seat-height conditions for the soleus and medial and lateral gastrocnemius. The effect of cadence on the EMG integral ratio differed between the two seat heights for all three muscles and between the muscles at respective seat heights. These data suggest that the direction of muscle action was not associated with the response of muscle activity to changes in cadence.  5.4 Additional Factors  Several factors potentially responsible for the effect of cadence on triceps surae EMG were not able to be tested in the current study. In this section, an explanation is given as to how these factors could influence muscle activation. 5.4.1 Pennation Angle  As pennation angle increases, the proportion of muscle fibre tension acting parallel to the muscle’s line of action decreases. To maintain a constant magnitude of muscle force with an increase in pennation angle, muscle excitation would be expected to increase. Several authors have demonstrated increases in pennation angle with decreases in muscle length (Fukunaga et al., 1997; Kawakami et al., 1998; Muraoka et al., 2001; Pasquet et al., 2005). Pennation angle could be a factor contributing to the differential effect of cadence on soleus and gastrocnemius EMG if the kinematic changes that occurred with increases in cadence led to a differential increase in pennation angle between the two muscles. 5.4.2 Moment Arm Length  The torque applied about a joint for a given magnitude of muscle tension is dependent on the length of the muscle’s moment arm. Determination of moment-arm length requires knowledge of the muscle’s attachment sites and of the joint’s instant centre of rotation. Since the soleus and gastrocnemius share a common insertion via the Achilles tendon, changes in moment arm length with changes in ankle angle will be similar between the two muscles. Changes in knee angle, however, will lead to changes in the moment-arm length of the gastrocnemius about the knee joint, but not of the soleus. To keep a constant magnitude of muscle torque about the knee with a varying moment- 50  arm length, the magnitude of muscle tension must also vary which would require changes in EMG. Differences in the effect of cadence on knee angle between the two seat heights would suggest that differences in both medial and lateral gastrocnemius moment-arm length about the knee joint between the two seat heights were also present, and could be a factor responsible for the differential effect of cadence on EMG. 5.4.3 Fibre Type  Fast glycolytic muscle fibres have faster contraction and relaxation times than slow oxidative fibres (Bottinelli et al., 1996). The medial and lateral gastrocnemii have been found to contain a greater proportion of fast glycolytic fibres than the soleus (see section 1.4.2) which may have resulted in a differential response to increases in muscle velocity. If the increased demands for muscle velocity with increases in cadence were greater than the velocity capacity of the slow oxidative fibres of the soleus, soleus may have been derecruited in favour of its synergists, the gastrocnemii.  5.5 Differential Functioning of the Gastrocnemii   Two main differences in response to seat-height and cadence manipulations were found between the medial and lateral gastrocnemius: 1) the two muscles exhibited different EMG profiles at any given cadence and seat height, and 2) the two muscles responded differently to the combined manipulations of cadence and seat height – medial gastrocnemius yielding a main effect of each factor, lateral gastrocnemius yielding a seat-by-cadence interaction. In their analysis of muscle activity during pedalling, Hug and Dorel (2007) noted that the medial and lateral gastrocnemii were active to varying degrees at different times in the pedal cycle. Differential activation of the gastrocnemii with changes in muscle length have been found previously by Cresswell et al. (1995) (first described in section 1.1.2) and by Shinohara et al. (2006). In both of these studies, surface EMG was recorded from the soleus, medial gastrocnemius and lateral gastrocnemius while participants were asked to generate isometric plantarflexor torque that matched various submaximal target levels. This procedure was carried out with the knee at varying degrees of flexion. In each study, one of the two gastrocnemii yielded an interaction between the magnitude of plantarflexor torque and muscle length while the other muscle yielded main effects of each factor without an interaction. These data, along with those of the current study, suggest that the medial and lateral gastrocnemius muscles undergo different patterns of excitation, and that conclusions arrived at through observation of one muscle should not be generalized to the other. 51   Differences in the reflexive innervation of the gastrocnemii could provide an explanation for the muscles’ differential EMG response. The medial and lateral gastrocnemii are innervated by the tibial nerve. The reflexive influences on activation of the soleus and entire gastrocnemius (generalized to both medial and lateral heads) have been well documented (Crone et al., 1987; Petersen et al., 1998, , 1999; Sinkjaer et al., 2000; Pyndt et al., 2003; Pyndt & Nielsen, 2003); however little research exists on the differential reflexive innervations of the medial and lateral gastrocnemius. The growing body of literature reporting differential changes in the activation patterns of these two muscles prompts research into their reflexive innervations.  Differences in the proximal attachment of the medial and lateral gastrocnemii likely resulted in differential changes in the contribution to knee flexor torque. In addition to this, differences in the architecture between the two muscles may explain some of the differences in activation. Lateral gastrocnemius has been found to have a greater fascicle length, that is, a greater number of sarcomeres in series, than the medial gastrocnemius, suggesting that the lateral gastrocnemius has the potential for greater displacement and velocity (Wickiewicz et al., 1983; Huijing, 1985). On the other hand, the medial gastrocnemius has been found to have a greater pennation angle, suggesting that it is at an advantage for the production of force. 5.6 Agonist-Antagonist Coactivation  Lombard’s Paradox (Lombard, 1903) describes the situation in which a bi-articular muscle is active at a time when the direction of the moment required at one of the joints that the muscle spans is opposite to that associated with the direction of muscle of shortening. Gregor et al. (1985) give the example of rising from a chair: it has been found that, at the time when a hip-extensor moment is required, the hamstrings and quadriceps muscle groups are active concurrently. The action of the hamstrings (and their synergists) overcomes the hip-flexor moment produced from the action of the bi-articular rectus femoris to yield a net extensor moment. The reverse occurs at the knee: the flexor moment from the hamstrings is overcome by the extensor moment from the rectus femoris (and its synergists) to yield a net knee extensor moment. The result is a combined movement of the hip and knee – extension and flexion, respectively – that would not be possible with activation of either the hamstrings or quadriceps muscle group, alone. However, the coactivation comes at a greater metabolic cost since each muscle group is working against the other at both joints.  Gregor et al. (1985) sought to determine whether Lombard’s paradox was also present in cycling, since the motion of the hip and knee during the propulsion phase of cycling (hip extension and knee extension) is similar to 52  that seen when rising from a chair. Four variables were compared across the pedal cycle: 1) the net EMG activity from the primary hip flexors/knee extensors (rectus femoris, vastus lateralis, vastus medialis), 2) the net EMG activity from the primary hip extensors/knee flexors (gluteus maximus, biceps femoris, semimembranosus), 3) the net moment about the hip, and 4) the net moment about the knee. In the first half of the propulsion phase, both muscle groups were active and both the hip and knee joint yielded a net extensor moment, consistent with the example of Lombard’s paradox seen in rising from a chair. However, in the second half of the propulsion phase, when the moments about the hip and knee were extensor and flexor, respectively, coactivation was reduced by a decrease in the activity of the hip flexor/knee extensor muscle group. This was interpreted by Gregor et al. (1985) as a means of avoiding the increases in metabolic demand seen with Lombard’s paradox by reducing the opposition of antagonist muscle groups.  Joint moments could not be calculated in the current study, since pedal forces were not recorded; however, the coordination of EMG activity between hip flexor/knee extensor and hip flexor/knee extensor muscle groups was consistent with that reported by Gregor et al. (1985), and did not change with increases in cadence or with a decrease in seat height. 5.7 Muscle Length Calculations   In this study, muscle-tendon unit lengths were estimated using pre-determined regression equations developed by Hawkins and Hull (1990) from measurements made on cadaveric specimens. The muscle-tendon length data were then used to draw conclusions regarding muscle function during cycling based on the muscles’ length-tension properties. To apply the estimates of muscle-tendon length in this way required the assumption of a certain degree of error, since the non-contractile, elastic elements of the muscle-tendon unit have been shown to change in length, independent of changes in length of the contractile elements. The magnitude of this error is based, in part, on the length of the elastic elements, specifically the muscles’ proximal and distal tendons, since a longer tendon has a greater capacity to stretch (Lichtwark et al., 2007). The non-contractile components of the medial gastrocnemius, including the proximal and distal (Achilles) tendons and the aponeurosis have been found to stretch up to 6% of their resting length during near-maximal voluntary isometric actions (Muramatsu et al., 2001; Maganaris & Paul, 2002; Lichtwark et al., 2007); the non-contractile elements of the tibialis anterior have been found to stretch up to 7% (Maganaris & Paul, 2002); and those of the vastus lateralis up to 4% (Muraoka et al., 2001). Despite its limitations, the Hawkins & Hull method (or similar methods described by Grieve et al. (1978) and 53  Visser et al. (1990)) have been used in several studies under the deliberate, erroneous assumption that the contractile and non-contractile elements of muscle cannot change in length independently (van Ingen Schenau et al., 1995; Sanderson et al., 2006). The impact of this error can be minimized by 1) avoiding conclusions based on changes in muscle length that are less than the possible changes in length of the non-contractile elements and 2) by ensuring low rates of muscular contraction. As the rate of rise of contraction increases, so does the likelihood that non- contractile elements will stretch to accommodate the shortening of the contractile components without a change in length of the entire muscle-tendon unit. The rates of rise of reaction forces on the foot from the pedal during cycling are less than those of ground reaction forces on the foot during weight-bearing activities such as walking and running (Patterson & Moreno, 1990; Winter, 1991). It follows that cycling is also associated with lower rates of muscle contraction. Thus, as a dynamic activity, cycling is better suited than walking and running for the estimation of muscle length from joint angle data. 5.8 Limitations  In this study, associations were made between dependent variables that yielded similar statistical results from a two-way repeated measures ANOVA. This procedure required the assumption that the associations between dependent measures did not differ between participants. The requirement of such an assumption was a weakness in this study’s statistical design. The manuscript based on this thesis that will be submitted to a peer-reviewed journal will be strengthened by the addition of correlation and regression analyses. 6.0 Conclusion   This study described the kinematic and electromyographic responses to changes in cadence and seat height during cycling. The most notable findings were:  1) The medial and lateral gastrocnemii responded differently to changes in cadence at each seat height. Kinematic differences could not explain the differential electromyographic response; a difference in muscle architecture seems the most likely cause. 2) Contrary to what the findings of past research would suggest (Cresswell et al., 1995; Tamaki et al., 1997; Miaki et al., 1999; Kennedy & Cresswell, 2001; Shinohara et al., 2006), the decrease in gastrocnemius activation in the low-seat condition was not associated with a decrease in gastrocnemius length, but may instead be due to changes in knee angle and vastus lateralis length. The findings 54  presented here are consistent with those presented by Arampatzis et al. (2006) and prompt further research on the neural pathways between the triceps surae and proprioceptors in the knee joint and knee extensor muscles. 3) It was evident that several factors, including muscle length, muscle velocity, ankle angle and the direction of muscle action, were not responsible for the differential effect of cadence on activation of the soleus and gastrocnemius. Preferential activation of the gastrocnemius during high-velocity movements has been demonstrated in humans (Nardone & Schieppati, 1988) and in cats (Smith et al., 1980), and suggests that muscle fibre composition is another potential factor mediating triceps surae activity.  This study expanded on previous research that investigated triceps surae activity during isometric actions with changes in lower limb kinematics by using a paradigm of dynamic activity, cycling. 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Eur J Appl Physiol Occup Physiol 33, 293-306.    64                          Appendix 1  Figures: Grand Means                                    65   66  67  68   69   70   71  72  73  74  75  76  77  78  79   80         81   82   83   84   85           86   87  88                          Appendix 2  Figures: Individual Participants                                    89   90  91   92  93   94  95  96  97  98  99  100  101  102  103  104  105  106  107  108  109  110  111  112   113  114   115  116  117  118  119  120  121  122                        Appendix 3  Behavioural Research Ethics Board Certificate of Approval    123  124                         Appendix 4  Participant Consent Form T H E  U N I V E R S I T Y  O F  B R I T I S H  C O L U M B I A  School of Human Kinetics  210 – 6081 University Boulevard  Vancouver, BC, Canada V6T 1Z1  Tel: (604) 822-3838  Fax: (604) 822-6842  Web: www.hkin.educ.ubc.ca  PARTICIPANT INFORMATION AND CONSENT FORM   Do changes in seat height modify the effect of cadence on activation of the triceps surae during cycling?   Principal Investigator: Dr. David J. Sanderson, Professor, School of Human Kinetics, University of British Columbia, phone: 604-822-4361  This study is part of a long-term project focused on the interaction between the soleus and gastrocnemius muscles of the lower leg during cycling exercise. Although these two muscles perform similar functions, each has different properties that can be revealed during cycling. Our goal is to explore some of these differences in this experiment. This project is unfunded research  Your participation is voluntary.  Before you decide whether or not you would like to take part in this study, it is important for you to understand what the research involves. This consent form will tell you about the study, why the research is being done, what will happen to you during the study and the possible benefits, risks and discomforts.  If you wish to participate, you will be asked to sign this form. If you do decide to take part in this study, you are still free to withdraw at any time and without giving any reasons for your decision.  If you do not wish to participate, you do not have to provide any reason for your decision not to participate nor will you lose the benefit of any medical care to which you are entitled or are presently receiving.  Please take time to read the following information carefully and to discuss it with your family, friends, and doctor before you decide.  You do not waive any of your legal rights to compensation by signing this consent form.  Purpose: The purpose of this study is to record the activity and length of the calf muscles as pedaling frequency and seat height are varied during cycling exercise of moderate intensity. 125  Procedures: The independent variables will be pedaling rate and seat height. The dependent variables will be the calculated lengths of the soleus and gastrocnemius muscles and the integrated electrical signals (EMG) from those two muscles.  When participants arrive at the lab, their height and weight will be measured. The two EMG recording sites will be shaved, lightly abraded and cleaned with isopropyl alcohol. The EMG electrodes will be fixed to the surface of the skin with double-sided adhesive pads. Reflective markers will be placed over the prominence of the hip, knee, ankle, heel and 5th metatarsal head. The bicycle seat height will be adjusted to suit the participant’s leg length.  Participants will perform five 3-minute trials at a power output of 200 W, considered a moderate intensity. The pedaling frequencies for the five trials will be 50, 65, 80, 95, 110 rpm with the order of pedaling frequencies being randomly selected. A minimum of three minutes of rest will be given between each trial. At the end of the five trials, the seat height will be adjusted and a second set of five trials will be performed.  During each trial, the participant’s heart rate will be monitored, and when it has reached a constant value, the EMG from the muscle groups will be recorded. At the same time as the recording of EMG, a video recording of the participant’s lower limbs will be made. The video tape will be digitized using the reflective markers to reveal the angular displacements of the knee and ankle joints. Data analysis will involve examining the dependent variables as a function of pedaling frequency at the two seat heights.  The total time for your involvement will be approximately 1½ hours.  Exclusion: If you meet any of the following criteria, you will be excluded from this study: • If you are outside the ages of 19-35 years • If you have no cycling experience • If you have any known neuromuscular disorder • If you have experienced a musculoskeletal injury to either of your lower limbs within the past 18 months.  Risks: The cycling intensity used in this study is considered to be moderate. Between each trial, there will be sufficient time to relax and recover. In the unlikely event that there is any discomfort we will stop the session. You may stop at any time. Depending on your current level of fitness, you may experience a small amount of muscle soreness for several days following this study.  Benefits: You will not receive any direct benefits by participating in this study.  Remuneration/Compensation: You will not receive any reimbursement for expenses, gifts-in-kind and/or payment by participating in this study. 126  Confidentiality: Any information regarding participant identification resulting from this study will be kept strictly confidential. All documents will be identified only by code number and kept in a locked filing cabinet. Data will be kept up to a maximum of five years and will only be accessible to Dr. Sanderson and Ryan Cawsey. Data files stored on computer will be labeled using code numbers on a computer in the UBC Biomechanics Laboratory. You will not be identified by name in any reports of the completed study.  Contact for information about the study: If you have any questions or require further information with respect to this study, you may contact Ryan Cawsey at 778-998-5304.  Consent: Your participation in this study is entirely voluntary and you may refuse to participate in or withdraw from the study at any time. If you are a student, your refusal to participate, or your withdrawal from the study will, in no way, affect your class standing. Your signature below indicates that you have received a copy of this consent form for your own records.  Your signature indicates that you consent to participate in this study.            ________________________________________________________________ Participant signature Date         ________________________________________________________________ Investigator signature Date 127 

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