In vivo muscle function vs speed I. Muscle strain in relation to length change of the muscle-tendon unit
1 Biological Sciences Department, California State Polytechnic University,
Pomona, CA 91768-4032, USA
2 Equine Research Center, California State Polytechnic University, Pomona,
CA 91768-4032, USA
3 Concord Field Station, Harvard University, Bedford, MA 01730,
USA
* Author for correspondence (e-mail: dfhoyt{at}csupomona.edu)
Accepted 22 December 2004
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Summary |
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Key words: Locomotion, quadruped, sonomicrometry, muscle
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Introduction |
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While these studies suggest that the direction and amount of strain in the
muscle fascicles can be uncoupled from changes in the length of the
muscle-tendon unit, there have been relatively few explicit examinations of
this during running, trotting or hopping. In one of the first studies of
muscle function using direct measurement of fascicle strain (sonomicrometry),
Griffiths (1991) found that
fascicles of the feline medial gastrocnemius shortened while the ankle flexed
(and the overall muscle-tendon unit was lengthened) during stance - a
phenomenon that Griffiths
(1991
) termed `mechanical
buffering'. A similar conclusion was reached for one of three cats, based on
the kinematic analysis of the function of the gastrocnemius
(Prilutsky et al., 1996b
).
This was a surprising demonstration because it had been assumed for years
(Walmsley et al., 1978
) that
the cat gastrocnemius, an ankle extensor, was undergoing an eccentric
(lengthening) contraction during this phase of the step cycle. Subsequent
sonomicrometry studies have not explicitly addressed this issue but their data
are relevant. In the studies that have quantified elastic storage in the
tendons of ankle extensors (Biewener et
al., 1998
; Roberts et al.,
1997
), the tendons stretch because the ankle joints are flexing
while the muscles shorten by small amounts. Ultrasonographic studies of human
walking have shown that the fascicles of the gastrocnemius function nearly
isometrically during lengthening of the muscle-tendon unit
(Fukunaga et al., 2001
) and a
recent ultrasonographic study of eccentric activity in human tibialis anterior
also showed that the fascicles behaved quasi-isometrically as the
muscle-tendon unit lengthened during ankle movement
(Reeves and Narici, 2003
).
Many of these studies are of ankle extensors with relatively long tendons
and short fascicles. Patterns of muscle strain in relation to joint angle
change are less consistent for more proximal muscles. The vastus lateralis (a
knee extensor) of the dog was found to shorten during the first half of stance
when the knee flexed (see fig. 4A in
Carrier et al., 1998). In
contrast, in a study of the vastus lateralis and biceps femoris (hip extensor)
of rats, Gillis and Biewener
(2001
) found a much closer
correspondence between joint motion and muscle strain. In the present study we
sought to extend these observations to the full range of normal trotting
speeds in horses, with the goal of comparing two muscles (forelimb triceps and
hindlimb vastus lateralis) in which one might expect to find substantial
differences in mechanical buffering.
During the stance phase of the stride in a trotting horse the elbow and the
knee exhibit different patterns of joint angle change that would be expected
to be the product of very different patterns of strain in their respective
extensor muscles. During stance, the elbow extends about 25° with a brief
period of flexion just before mid-stance
(Back et al., 1995a). Thus, one
would expect that elbow extensors would shorten throughout stance with,
possibly, a brief period of eccentric activity near mid stance. The knee
exhibits two phases during which flexion is followed by extension
(Back et al., 1995b
;
Hoyt et al., 2002
). The knee
extensors are active during most of stance
(Tokuriki and Aoki, 1995
) as
they must stabilize the knee in order to support the animal's body weight.
Given these observations, there would seem to be at least two possible
patterns of muscle strain in the vastus. If muscle strain closely follows the
joint kinematics, then the muscle would undergo two cycles of eccentric
activity followed by concentric activity. These `stretch-shortening' cycles
would be expected to increase force production (Ettema,
1996
,
2001
;
Hof et al., 1983
).
Alternatively, if the muscle's series elasticity is sufficiently compliant,
then during the apparently eccentric phases the muscle might remain isometric
or even shorten, as Griffiths
(1991
) observed in the cat
medial gastrocnemius. In the present study we measured muscle strain in the
lateral triceps (an elbow extensor) and in the vastus lateralis (a knee
extensor). Lacking specific information on compliance of their series
elasticity, we hypothesized that muscle strain would track joint excursion in
both muscles: the triceps shortening throughout stance and the vastus
exhibiting two cycles of eccentric activity followed by concentric
activity.
We were also interested in testing the hypothesis that the amount of strain
in both muscles would increase with speed. We expected strain to increase
because step length increases with speed
(Farley et al., 1993;
Hoyt et al., 2000
) and
increased step length should require increased range of motion in leg joints,
a phenomenon previously observed in the equine knee
(Hoyt et al., 2002
). Because
time of contact decreases with speed in a trot
(Dutto et al., 2004b
;
Hoyt et al., 2000
) we also
expected to find that strain rate would increase with speed, even if the
amount of strain did not. Previous studies of in vivo muscle function
in quadrupeds (Carrier et al.,
1998
; Gillis and Biewener,
2001
; Gregersen et al.,
1998
) have also studied changes in muscle function with speed.
However, in these studies the increased speed was accompanied by changes in
gait, which may confound the results from speed alone.
Muscle strain can be active or passive. In order to focus on strain when
the muscle was active, EMG (electromyographic) activity was measured relative
to the time the hoof was on the ground. These data also allowed us to
determine the duty factor, the proportion of the stride during which the limb
is in contact with ground. Duty factor is related to peak forces
(Alexander et al., 1979) and,
therefore to the total volume of muscle active
(Taylor, 1994
). Because we
hypothesized that strain rate would increase with speed, we also predicted
that the volume of actively recruited muscle, as indicated by the integrated
EMG (IEMG) would increase with speed.
Because a preliminary study revealed unexpected variability between individuals in the strain pattern of the vastus, we implanted sonomicrometry crystals in approximately the same locations within the vastus lateralis muscles of four horses in two consecutive summers and compared the patterns of strain observed throughout the stride cycle.
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Materials and methods |
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Surgical procedures
Surgery was done on standing, sedated horses. On the morning of the
surgery, each animal was catheterized using a 16-gauge catheter to facilitate
the administration of drugs. An initial dose of xylazine (xylazine HCl;
Fermenta Animal Health Co., Kansas City, MO, USA) was given (1.1 mg
kg-1) to tranquilize the horse. A combination of butorphanol
tartrate (Fort Dodge Animal Health, Fort Dodge, IA, USA; 0.1 mg
kg-1) and detomidine hydrochloride (Pfizer Animal Health, Exton,
PA, USA; 20-40 µg kg-1) was administered intravenously at the
beginning of the surgery and throughout the surgery as needed to maintain
sedation. The locations of the lateral triceps (M. triceps brachii caput
laterale) and vastus (M. vastus lateralis) were determined by palpation in the
right limbs. The approach to the lateral head of the triceps was identified by
palpation of the lateral epicondyle, the deltoid tuberosity of the humerus,
and the olecranon of the ulna. The vastus was identified using the landmarks
of the greater and lesser trochanter and the lateral ridge of the trochlea of
the femur. Anatomical locations were studied on several cadavers prior to
surgery and anatomical validation of sonomicrometer crystal placement was done
on three horses not part of this study that were euthanized for other medical
conditions not related to musculoskeletal dysfunction.
Lidocaine HCl (Pro Labs Ltd., St Joseph, MO, USA) was administered subcutaneously at the incision sites and followed by a 7 cm long incision in the skin. The fascia was exposed by removing subcutaneous fat and, in the case of the triceps, incision through the omobrachialis muscle. Two small stab incisions were made into the fascia, approximately 10-15 mm away from each other in a line parallel to the muscle fiber orientation. For each muscle, one pair of 2 mm omni-directional, spherical, piezoelectric crystals (Tack crystals, Sonometrics Corporation, London, Ontario, Canada) was implanted to measure changes in muscle fiber length. Crystals were implanted 1 cm deep into the muscle using a polyethylene introducer. The crystals were anchored to muscle fascia using 0 silk suture, and a tension relief loop. The sonomicrometry output was sampled at 463 Hz using Sonometrics System Software and then output to the A/D card (PCI 1200, National Instruments, Austin, TX, USA), which acquired the data at 3704 Hz using LabVIEW® software (National Instruments, Austin, TX, USA) running on an IBM-compatible personal computer. The higher frequency was necessary for the EMG signals that were being acquired by the same A/D system, and the resulting multiple records of the sonomicrometry signal were removed during data processing.
Electromyography electrodes, made of multi-stranded, Teflon-coated,
stainless steel wire (AS636, Cooner Wire, Chatsworth, CA, USA) were inserted
by a sew-through technique (Carrier,
1996) 1 cm away from, and parallel to the sonomicrometry crystals.
The modification made to the procedure of Carrier was the absence of a
proximal silastic patch to anchor the EMG wire to the fascia. This reduced
muscle tearing that might have resulted from the presence of two fixed
anchors. Each of the two EMG electrodes had 7 mm of bared wire and the two
segments were separated by 1.5 cm. The EMG signal was amplified (1000-10 000,
depending on signal strength) and filtered (60 Hz notch and 100-1000 Hz
bandpass) with a Grass model P511K preamplifier (Quincy, MA, USA), and sampled
at 3704 Hz by the A/D system. A ground wire was implanted subcutaneously into
the dorsal aspect of the horse's sacral region. Following implantation, all
incisions were loosely sutured with 0 silk. Flunixin meglumine
(Schering-Plough Animal Health Corp., Union, NJ, USA; 20-40 µg
kg-1) was administered post-surgically to reduce inflammation and
associated pain.
Data collection
After the surgery was completed, the animal recovered for at least 90 min
and was allowed to drink water and graze freely. After this time period, a
lameness examination was conducted by a veterinarian to ensure soundness in
all four limbs and provide a subjective measure of any residual effects of
sedation or surgery. An objective measure of surgical treatment and sedation
was also used: measurements of heart rate and manually calculated stride
frequencies during data collection were compared to pre-surgical measurements
taken during exercise bouts.
A surcingle was placed around the horse slightly posterior to the scapula and was used to secure the wire connections from the implants and accelerometers. A biaxial accelerometer (±50 g, CXL25M2, Crossbow Technology, Incorporated, San Jose, CA, USA) was taped on the lateral aspect of the hoof of the right fore- and hindlimbs to record hoof contact and break-over (the end of stance when the hoof leaves the treadmill). All accelerometer data were collected at 3704 Hz and analyzed using the LabVIEW software. The validity of the accelerometer as a measure of hoof contact and break-over was assessed by comparing accelerometer tracings with simultaneous synchronized high-speed video.
Each horse was run on a high-speed treadmill under two conditions: on the
level and up a 10% incline. The data collected on the level are the focus of
this paper and those collected on the incline are the focus of a companion
paper (Wickler et al., 2005).
Once on the treadmill, the horse was warmed up at a walk (1.7 m
s-1) for 3 min and a trot (3.0 m s-1) for 5 min. Horses
were run under each condition at speeds from 2.5 to 4.5 m s-1 in
0.25 m s-1 increments. The conditions and speeds were randomly
ordered. Horses were brought up to speed, and after 45 s at speed, data were
collected. All 18 experimental trials (nine of each speed at 0% and at 10%
incline) were run in succession, with a 30 min break after the first nine.
Following the last trial (defined as a combination of speed and condition),
the horse was again sedated and sonomicrometry crystals and EMG wires were
removed. Following this surgery, the animal received phenylbutazone (Pro Labs
Ltd., St Joseph, MO, USA; 11-22 mg kg-1) and was returned to an
open paddock. The horses were hand-walked daily for 1 week, following which
sutures were removed.
Kinematic data
For the study of knee and elbow kinematics, reflective markers (Peak
Performance Technologies, Englewood, CO, USA) were glued to the skin on the
lateral side of the right fore- and hindlimbs of the horses using standard,
palpable positions for these joints (Back
et al., 1993). The markers were 2.5 cm, lightweight, plastic
spheres covered with corner-cubed reflective tape. The joint angles were
calculated so that a decrease in joint angle indicates a flexing joint. The
horses were filmed at 125 Hz using a Model PCI Motion Scope® camera
(Redlake Camera Corp., Morgan Hill, CA, USA). The camera was placed
approximately 8.5 m away from the treadmill. A linear calibration was
performed daily. Five consecutive strides were recorded and digitized using
Motus® software (Peak Performance Technologies, Englewood, CO, USA) during
each sonomicrometry trial. The angular data were smoothed using a cubic spline
filter, normalized for time using a cubic spline interpolation and five
strides for each horse, speed and condition were averaged using the trial
averaging feature of Motus. These data were used to interpret the muscle
strain data (see below) and to determine mean joint angle of the knee and
elbow at first hoof contact, break-over, mid-stance and maximum extension
(elbow) or maximum flexion (knee).
Strain estimated from kinematics
The change in length of the muscle-tendon unit at 3.5 m s-1 was
estimated from the kinematic data, the length and pennation angle of the
muscle fibers, and the moment arm of the muscles. This length change was
converted to strain by dividing by the observed length of the muscle fascicles
and compared with the muscle strain observed during the same period of time.
Kinematic data from the elbow were not corrected for skin displacement because
the errors are very small (Back et al.,
1995a). Kinematic data for the knee were corrected for skin
displacement (van Weeren et al.,
1992
) at all four points used in calculating knee angle.
Individual horse's kinematic records were used to determine the range of
motion of the elbow and knee during each of the phases used for analysis of
muscle strain.
The moment arms of the equine triceps and vastus were determined from cadavers of four, similar-sized horses (based on body mass and linear dimensions of metacarpal and metatarsal bones). Muscles around the joints, other than the muscle of interest, were removed to permit the movement of the joint through the range of motion found in our kinematic studies. The distance along the line of action of the tendon was measured with the joint held at five different angles spanning the range of motion. Joint angles were measured to the nearest 0.5° with a protractor. For the forelimb, measurements were made from a pin placed into the most prominent palpable aspect of the deltoid tuberosity to a pin placed into the middle of the olecranon at a point approximately 4 cm from the posterior margin. For the hindlimb, measurements were made from a pin placed in the third trochanter on the midline of the femur to a pin placed in the center of the patella. The moment arm was calculated as the slope of a least-squares regression of the length of the muscle on the angle of the joint (in radians). Calculating moment arm in this manner means that a single average value was used at all joint angles.
The observed motion of the joint multiplied by the average moment arm was
used to predict the length change of the muscle-tendon unit during each phase
of muscle activity. This length change was converted to strain by dividing by
the average length of the muscle fascicles observed in three cadavers. The
equine vastus arises from the femur and inserts on the lateral margin of the
rectus femoris where there is a thick aponeurosis that effectively serves as
the tendon for the vastus. The angle the vastus muscle fibers made with this
aponeurosis was measured on three cadavers with a protractor and averaged
60.3°. For each of the four individual horses used in the sonomicrometry
measurements, the observed motion of the joint multiplied by the average
moment arm from the cadavers was used to predict the length change of the
tendon during each phase of muscle activity. The length change of the tendon,
corrected for the pennation angle, yielded the predicted length change of the
muscle fascicles (Gans and DeVree,
1987). This length change was converted to strain by dividing by
the average length of the muscle fibers measured on the three cadavers.
Data processing
Initial hoof contact, break-over and subsequent hoof contact were
determined using the record from the accelerometer, and from these were
calculated duration of stance phase (tc=time of contact)
and duration of swing phase. All other stride parameters were derived from
these measurements and speed. The EMG and sonomicrometry records were
subdivided into into individual strides using the simultaneously collected
accelerometry record.
Muscle length changes (and velocities of shortening) were analyzed only for
the time of contact because of its central role in determining metabolic cost
(Kram and Taylor, 1990),
although recent work identifies a significant energetic cost associated with
the swing phase (Marsh et al.,
2004
). The voltage output of the sonomicrometer was converted to
length using an empirically derived equation relating voltage to the actual
distance between the crystals, measured in water with the crystals attached to
a micrometer, similar to the procedure of Gillis and Biewener
(2001
). Muscle lengths from
the sonomicrometry records of the stance phase were smoothed using a median
pass filter that scanned through the data, substituting for each consecutive
datum, the median of the seven consecutive values centered on that datum (the
datum, the three preceding and three succeeding values). All records of muscle
lengths were normalized by dividing each value by resting muscle length. The
measurements of resting muscle lengths were taken when the animal was standing
square on the treadmill (metacarpals and metatarsals perpendicular to the
treadmill's surface). In order to calculate strain, the normalized muscle
length records of ten consecutive individual strides were averaged after being
temporally normalized to 100% of time of contact using a cubic spline
interpolation.
Strain patterns were not simple monotonic changes (particularly for the vastus), so that a simple calculation of the strain based on the difference in muscle length between the start of stance and the end of muscle activity would obscure considerable information about the relation between muscle strain and limb motion. Therefore the average normalized muscle length data were divided into phases based upon the kinematics of the appropriate joint. This was possible because the joint kinematics were determined from five strides that were a sub-set of the ten strides averaged for muscle length. The net strain (change in muscle length, with shortening being negative) occurring during each phase was determined for each animal and trial. Stain rate (muscle lengths per second) during each phase was determined from the net strain and the duration of the phase in that trial.
Electromyography records were filtered using a second order low pass filter (1000 Hz), rectified and integrated, and analyzed for: (1) when the EMG started relative to hoof contact, (2) total duration of the EMG signal (including if it started before stance), (3) the length of the signal (only during stance) as a percentage of tc, and (4) the integrated EMG during stance.
Reproducibility of sonomicrometry data
In an attempt to assess the reproducibility of inter-individual differences
in sonomicrometry data, we compared records of changes in muscle length over
the stride period from crystals implanted for a single day in the same horses
in two consecutive summers. For this analysis we used a subset of the data
reported here (limited to data collected on the level and four speeds) which
were the only data available that corresponded to a set of data obtained the
previous summer from the same horses. All relevant details of the experiments
were the same in both summers and the data were processed in the manner
described above except that all muscle length data from both summers were
normalized by dividing by the average length of the muscle during the stance
phase while trotting at 3.0 m s-1 because we did not have standing
square values for the first summer. Time was normalized by dividing by the
stride period. Data were analyzed from three horses, trotting at four speeds
(2.5, 3.0, 3.5, 4.0 m s-1) on the level. Values for every 0.5% of
the stride period were obtained by cubic spline interpolation. Reproducibility
of the data was assessed by calculating the variance ratio
(Hershler and Milner, 1978),
which is the ratio of the average variance between corresponding points in
different strides to the total variance of the entire data set.
Statistics
A two-way analysis of variance with repeated measures was run on all data
using SuperANOVA® software (Abacus Concepts Inc., Berkeley, CA, USA) with
significance set at P<0.05. The two variables tested were speed
and condition, and the four horses were used as the repeated measure. The data
analyzed were mean values for each horse, based on five strides for the
kinematic parameters and ten strides for all other parameters. All means are
presented ± S.E.M.
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Results |
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Our analysis of strain in the vastus (Fig. 3) was based upon the kinematics of the knee (Fig. 2) and the EMG and muscle length changes of the vastus (Table 2). The pattern of kinematics and muscle length in the vastus was more complex than the patterns recorded for the triceps. In part this was because there was a brief period of muscle lengthening (phase 2; Fig. 2Bi) in two of the animals that was not evident in the other two (Fig. 2Bii). The start of phase 2 was defined by the start of muscle lengthening and not kinematics, the only phase in either muscle that was not defined exclusively on the basis of kinematics.
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Phase 1 in the vastus lasted until 18% of stance at all speeds
(P=0.607). The knee flexed but the muscle shortened by
10.0±0.3%. Because phase 2 was only observed at all speeds in one
animal (where it averaged about 6% eccentric strain), we have omitted it from
the analysis. On average, phase 3 lasted from 18% to 36% of stance. The end of
this phase occurred progressively later (P=0.041) at higher speeds
(increasing from 29% to 39%) so that the relative duration of phase 3
increased (P=0.001). However, owing to the decreased time of contact,
there was no change (P=0.52) in the absolute temporal duration of
this phase. Seen in all four horses, phase 3 was a period of concentric
activity during which the knee extended and the vastus shortened by an average
of 3.0±0.6%. Strain increased (P=0.016) from 2.8% to 4.6% as
speed increased from 2.5 to 4.5 m s-1
(Fig. 3). Phase 4 was a period
of knee flexion that lasted, on average, from 36% until 61% of stance. The
duration of phase 4 was relatively consistent (15% of stance) at all
speeds because the end of the phase increased (P<0.001) with speed
from 55% to 65% of stance. This was a period in which the muscle underwent
minimal strain (Fig. 3),
shortening by 2.0±0.1%. The EMG was active until 59% of stance, very
near the end of phase 4. Phase 5 lasted until an average of 84% of stance and
was another period of concentric activity when the knee extended and the
muscle shortened by an average of 3.1±0.4%. Both the relative
(P=0.024) and absolute (P=0.002) duration of phase 5
decreased with speed. During this phase, strain in the vastus
(Fig. 3) increased from 1.6% to
4.1% with speed (P=0.047). Phase 5 occurred after the end of EMG
activity and the average time from the end of EMG activity until the end of
phase 5 was 0.069 s.
Muscle strain rate
In the triceps the strain rate (Table
1) during phase 2 averaged -0.112±0.052 fascicle lengths
s-1 and the effect of speed was nearly significant
(P=0.079). During phase 3 in the triceps
(Fig. 3) the strain rate
increased with speed (P=0.007). In the vastus, strain rate
(Fig. 3) remained constant
(Table 2) in phases 1
(P=0.306) and 4 (P=0.978) and increased with speed during
phases 3 (P=0.034) and 5 (P=0.005).
Electromyograms
The start of EMG activity in the triceps
(Fig. 4A) preceded foot contact
(termed `phase advance') by an average of 0.036±0.001 s and did not
change with speed (P=0.746). In contrast, in the vastus, phase
advance increased nearly linearly with speed (P<0.0001): at 2.5 m
s-1 the onset of EMG activity coincided with the start of stance
(phase advance = 0.003±0.007) but at 4.5 m s-1 the EMGs
started 0.069±0.018 s before stance. The total duration of the EMG
decreased with speed (P<0.001) in the triceps but stance duration
also decreased with speed. Hence, the percentage of time the EMG was on during
stance (Fig. 4B) did not change
with speed in the triceps and averaged 40±1% of stance. In the vastus,
the total duration of the EMG did not change with speed (P=0.274).
However, because of the increase in vastus phase advance, there was a decrease
in the time during stance when the muscle was active that paralleled the
decrease in time of stance. As a result, the percentage of time that the EMG
was on during stance (Fig. 4B)
also did not change with speed in the vastus (P=0.084) and averaged
59±2%. EMG duration in the triceps was shorter than in the vastus
(P=0.026).
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Integrated EMGs (Fig. 4C),
normalized to the highest value for that muscle in each horse when trotting up
an incline (Wickler et al.,
2005), and including only the portion that occurred during stance
when muscle force would be supporting the animal's weight, increased with
speed in the triceps (P=0.024) and the vastus (P=0.0001).
Average EMG phase advance was not different between the two muscles
(P=0.622), averaging 0.036±0.001 s in triceps and
0.042±0.005 s invastus. However, this comparison was complicated by the
fact that phase advance was nearly constant at 0.036 s in triceps and
increased from 0 s to 0.069 s in the vastus.
Kinematics
The patterns and magnitudes of the joint kinematics of the fore- and
hindlimbs during stance were similar to those reported by Back (1995a,b). The
angle of the elbow at contact (107±1.5°) did not vary with speed
(P=0.373). The elbow extended until very near the end of stance
(97%) and the angle at maximum extension of the elbow
(146±1.2°) was not different with speed (P=0.540).
Consequently, the range of motion of the elbow from contact to maximum
extension (39±0.8°) did not differ with speed (P=0.754).
The timing of maximum extension of the elbow in the stance phase increased
with speed (P=0.0001) from an average of 91.2±1.5% of stance
to 98.7±3.9% of stance. Following maximum extension, the elbow flexed
slightly but this was after the end of activity in the triceps.
The angle of the knee at contact (153±1.0°) did not vary with
speed (P=0.837). The knee underwent two cycles of flexion and
extension during stance (see descriptions of the phases of muscle activity,
above, for the timing of these events). The angle of the knee at the end of
stance (123±1.0°) was not affected by speed (P=0.279), and
consequently, the range of motion of the knee from contact to end of stance
(30±0.5°) was not affected by speed (P=0.384). All of the
angles reported above were not corrected for skin displacement errors so that
they could be compared with data in a companion paper
(Wickler et al., 2005) dealing
with trotting up an incline - a condition for which corrections for skin
displacement errors have not been reported. However, in general, correcting
for skin displacement did not change the timing of kinematic events but
decreased the range of motion of the knee by about 1.0°.
Stride parameters
Although stride period, stride frequency and stride length did not differ
between the limbs, the two components of stride period (swing and stance) were
different. As expected, stride period decreased with speed (P=0.0001,
Fig. 5A), resulting in a linear
increase in stride frequency with speed (P=0.0001). Stride length
also increased with speed (P=0.0001). Whereas swing duration
(Fig. 5B) did not vary with
speed in the forelimb (P=0.500), it decreased monotonically with
speed in the hindlimb (P=0.0001) and averaged 10% shorter in the
forelimb than in the hindlimb (P=0.001). Stance duration
(Fig. 5C), also called time of
contact (tc), decreased with speed in both limbs
(P=0.0001), but forelimb stance duration averaged 12% longer
(P=0.046). Step length (Lc), the distance the
body moves during time of contact (the product of stance duration and speed),
increased with speed (Fig. 5D)
in both limbs (P=0.001), but the longer time of contact in the
forelimb produced a step length that was on average 11% larger
(P=0.046). Because swing time did not decrease in the forelimb, time
of contact decreased more with speed than in the hindlimb and the difference
in time of contact between the two limbs decreased with speed
(P=0.001) from an average of 0.048 s at 2.5 m s-1 to 0.021
s at 4.5 m s-1. The inverse of time of contact
(1/tc) increased linearly with speed
(Fig. 5E) in both limbs
(P=0.0001) but it was 10% lower in the forelimb (P=0.044).
Duty factor, the proportion of time the foot is in contact with the ground
during the entire stride (duration of stance divided by stride period),
decreased with speed (Fig. 5F)
in both limbs (P=0.0001), but the combination of a longer time of
contact for the forelimb with identical stride periods produced a 12% larger
duty factor (P=0.015).
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Strain estimated from kinematics
The estimates of muscle moment arms (taken as the slope of the linear
regression of muscle length on joint angle in radians) seem fairly reliable
because the two variables were highly correlated (the average
r2 was 0.98 and the lowest was 0.96). For each of the
phases, the length change of the fascicles at 3.5 m s-1 was
predicted from the product of the observed range of motion, as determined from
the kinematics, and the empirically determined moment arm (triceps:
5.39±0.26 cm; vastus: 6.46±0.24 cm). This was converted to
strain by dividing by the observed length of the muscle fascicles (triceps:
21.7±0.6 cm; vastus: 20.0±0.8 cm) and compared with the values
observed by sonomicrometry (triceps, Table
3; vastus, Table
4). In the triceps the observed length change is the opposite of
the prediction during phases 1 and 2. The difference during phase 1 is
inconsistent: prediction of -6.4% (shortening because the joint is extending)
compared with +3.0% (lengthening) observed. During phase 2 the differences are
not nearly as large (e.g. +2.0% predicted because the joint is flexing and
-0.7% observed). During phase 3 the observed and predicted shortening are very
similar (-9.7% predicted and -11.3% observed).
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|
The differences are more pronounced in the vastus (Table 4). During the two concentric phases in the vastus (phases 3 and 5), the amount of strain observed via sonomicrometry is about 3.8 times the amount predicted from kinematics (-1.9% predicted and -7.2% observed). However, during phase 4 (the expected eccentric phase near mid-stance) the muscle was predicted to lengthen by 3.2% but it actually shortened by 3.7%. In both muscles the largest difference between predicted and observed values occurred during phase 1 (triceps: -6.4% vs +3.0; vastus: +2.9% vs -9.4%; predicted and observed, respectively).
Reliability of sonomicrometry data
The variance ratio was used to assess repeatability of the temporal muscle
waveform data obtained with sonomicrometry for full strides. It is the ratio
of the average variance at regular points along the waveforms for consecutive
strides to the total variance of all points in these strides. This ratio is
zero if there is perfect repeatability of the strides since each point is
exactly the same for each stride, i.e., the variance at each point is zero.
Likewise, the ratio is one if the waveforms are completely random (no
repeatability) since the average for variances at each point should approach
the variance of all points. Normalized vastus muscle length data from two
consecutive years in two different horses are presented in
Fig. 6. These two figures
represent the range of variance ratios obtained in this analysis. The
reliability (Table 5), which is
the variance ratio for ten consecutive strides in one trial averaged 0.029
(range 0.009-0.08). Constancy (Table
6), which is the variance ratio over the 2 years, averaged 0.149
(0.027-0.250). An ANOVA indicated that reliability and constancy did not vary
significantly between speeds or years.
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Discussion |
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While the triceps did shorten during stance, perhaps the most interesting result of this study was the fact that, in both extensor muscles, periods of joint flexion were not associated with eccentric activity. The strain pattern of the triceps, a parallel-fibered extensor of the elbow, can be considered to be composed of three phases on the basis of elbow kinematics (Fig. 2). The second of these phases is an eccentric phase during which the elbow flexes by 2.3°. The triceps would be expected to lengthen by 1%, but it actually shortened 0.7% (values that are significantly different). The vastus, a unipennate extensor of the knee exhibits a complex strain pattern (Fig. 2) including a phase of flexion (phase 4) of 11.5° during which the muscle might be expected to lengthen by 3.2% but actually shortened by 3.7%. In both muscles, these phases of activity occurred when the muscle was unambiguously active, starting at about 20% of stance and ending just prior to the end of EMG activity (Tables 1 and 2).
For neither muscle does it seem likely that the difference between observed
and predicted strain during the eccentric phase can be completely explained by
strain in an external tendon. First, the triceps has no external tendon to
account for this difference. Secondly, the amount of calculated strain in the
tendon across the knee seems unreasonable. For example, during the initial
phase of activity and knee flexion (phase 1), when the muscle was predicted to
lengthen by 2.9% it actually shortened by 9.4%. The total difference (12.3%)
amounts to about 2.5 cm in the vastus, which has an average fascicle length of
about 20 cm. The total length of the middle patellar tendon is about 12 cm, so
if all of this difference were absorbed by the muscle's tendon, the strain
would be about 20% - far greater than any tendon can sustain without breaking
(Shadwick, 1990). Some of the
difference between predicted and observed strain may be due to the fact that
measurements were made on different animals (tendon lengths, muscle moments
and fascicle lengths were calculated from cadavers). An additional factor that
contributes to some uncertainty in the present prediction of strain from
kinematics and morphometrics is the fact that there is a range of fascicle
lengths in both muscles. However, some confidence in these calculations is
supported by the observation that in both muscles there were phases of
concentric activity (triceps phase 3; vastus phases 3 and 5) when the muscle
shortened by values that are not significantly different than predicted from
kinematics (Tables 3 and
4).
If the external tendon cannot account for all of the difference between
length changes of the fascicles and the length change of the muscle-tendon
unit, then some of the difference must be due to strain in the aponeuroses
within the muscle. There is strong evidence that aponeuroses are an important
component of a muscle's series elasticity
(Fukashiro et al., 1995;
Huijing and Ettema, 1988
;
Maganaris et al., 2001
;
Monti et al., 2003
;
Scott and Loeb, 1995
) but
there is conflicting evidence as to whether there are differences in the
amount of strain in tendons and aponeuroses and whether there is heterogeneity
of strain within the aponeuroses (Finni et
al., 2003
; Maganaris,
2002
; Maganaris and Paul,
2000b
; Muramatsu et al.,
2001
). Maximum reported strain values for tendons are usually
around 5-7% and values for aponeuroses range from 2% to 14.3%
(Muramatsu et al., 2001
).
However, in some cases the tendon is reported to be more compliant than the
aponeurosis (Magnusson et al.,
2003
). While there have been no direct studies of series
elasticity in the equine vastus lateralis, there is substantial strain in the
elastic elements of the human vastus lateralis
(Bojsen-Moller et al., 2003
;
Fukunaga et al., 2002
;
Ichinose et al., 2000
;
Muraoka et al., 2001
) and at
least one study (Finni and Komi,
2002
) supports the idea that, in muscles in general, the
aponeurosis may be more compliant than the tendon
(Kawakami and Lieber, 2000
).
In the absence of data on the compliance of the series elasticity of the
equine vastus lateralis, we cannot unambiguously account for the differences
in observed and predicted strain.
While both muscles probably store some strain energy in series elastic elements such as the aponeurosis, it is striking that this energy does not appear to be recovered through subsequent extension of the joints. The end of vastus activity is estimated to be at about 75% of stance and the timing of maximum flexion changes from 63% of stance at 2.5 m s-1 to 72% of stance at 4.5 m s-1, indicating that the joint extends for only a short period of time at the end of muscle activity when energy stored in stretched elastic structures might be recovered. During this time the knee extended only slightly (an average of 4°), suggesting that some elastic energy may have been recovered from the stretched patellar tendon and the muscle's aponeurosis.
These data are consistent with the hypothesized triceps strain pattern but
not the hypothesized vastus strain pattern. The total strain in the triceps
(Table 1) during phases 2 and 3
(-10.6%) is very similar to that observed in the vastus
(Table 2) during phases 3
through 5 (-8.1%), despite the fact that the two joints have very different
actions: the elbow extending and the knee flexing during limb support. The
fact that the vastus shortens even though the knee flexes during most of
stance is a good example of the importance of series elasticity as a
mechanical buffer (i.e. protecting against potentially damaging rapid stretch
of muscle fibers to longer lengths) during potentially eccentric activity
(Fukunaga et al., 2001;
Griffiths, 1991
;
Reeves and Narici, 2003
).
The results of this study can also be used to test Taylor's `economical
force hypothesis' (Taylor,
1994) that the muscles responsible for supporting the body during
the stance phase of the stride function largely in an isometric fashion.
Although both muscles have complex strain patterns, our results generally
support Taylor's hypothesis. The net amount of strain observed in the horse
triceps (-10.6%) and vastus (-8.1%) is only slightly more than the
6%
shortening strain reported for more distal, pennate muscles in other species.
The lateral gastrocnemius of wild turkeys
(Roberts et al., 1997
)
shortened by 6.6±1.9% of resting length. In tammar wallabies
(Biewener et al., 1998
), the
plantaris shortened less than 2%, and the lateral gastrocnemius less than
6%.
The strain in the horse vastus (-8.1%) is somewhat less than the amount of
strain reported in extensors of other species. In trotting dogs, the amount of
shortening in the vastus lateralis (Carrier
et al., 1998) ranged from 8.7% to 20.1%, whereas the
semimembranosus shortened by 4.5% to 6.6%. The biceps femoris (a hip extensor)
shortened by an average of 27% during trotting in rats
(Gillis and Biewener, 2001
),
whereas the vastus lateralis lengthened by 8% to 16% while electrically active
during stance. It is remarkable that, although the kinematics of the knee are
relatively similar in the rat, dog and horse, the strain pattern in the rat
vastus differs substantially. In all three species, the knee flexed during at
least the first half of stance while the muscle was electrically active. In
the rat, the vastus lengthened continuously by about 10% during this period of
time.
Strain vs speed
The results of this study support, in part, our third hypothesis that
muscle strain should increase with speed. Although concentric (shortening)
strain in the triceps did not change with speed, there was a significant
increase during both concentric phases in the vastus. For the triceps this was
not surprising in view of the fact that the range of motion of the elbow, from
contact to maximum extension, did not change with speed. However, the range of
motion of the knee, from contact to maximum flexion, did increase with speed
(by 5°), and as a knee extensor, the vastus might have been expected to
shorten less at higher speeds. In the tammar wallaby
(Biewener et al., 1998)
fascicle strain within the plantaris and gastrocnemius muscles did not
increase with hopping speed between 2.5 and 6.0 m s-1. In the
report on the turkey (Roberts, 1997) it is not clear whether the increased
muscle work with speed (fig. 3 in Roberts, 1997) was due to increased strain,
increased force development, or both. None of the other in vivo
studies of muscle function (Carrier et al.,
1998
; Gillis and Biewener,
2001
; Gregersen et al.,
1998
) examined the effect of speed within a single gait.
Strain rate
Our results support the fourth hypothesis that strain rate would increase
with speed. In both muscles, strain rate increased with speed during the
concentric phases: triceps phase 3 (Fig.
3) and vastus phases 3 and 5
(Fig. 3). Anincreased muscle
strain rate with speed might result from the following three conditions: the
same amount of strain occurring in less time, more strain occurring in the
same time, or a combination of more strain and less time. In the triceps the
amount of strain did not change with speed but the absolute duration of phase
3 did decrease. In the vastus there was increased strain with speed during
both phases. The absolute duration of phase 3 did not change with speed but it
decreased for phase 5. Thus, considering the three concentric phases in these
two muscles, the increased strain rate resulted from all three combinations of
the two variables. The fascicle-shortening rate observed during phase 3 in the
triceps (0.8 lengths s-1) was much less than the long head of
triceps in the dog (1.4 to 3.6 lengths s-1;
Gregersen et al., 1998). The
horse vastus strain rates during phases 3 and 5 (0.6 to 0.8 lengths
s-1) were also slower than those reported for the dog vastus
(1.26-2.62 lengths s-1; Carrier
et al., 1998
), but similar to those of the dog semimembranosus
(0.66-1.07 lengths s-1;
Gregersen et al., 1998
).
Because the strain rates reported in previous sonometric studies were obtained
from longer, sustained periods of concentric activity, rather than from
shorter, kinematically defined, components of stance, the values may not be
directly comparable. Gillis and Biewener
(2001
) reported that during
trotting the strain rate of the rat biceps femoris increased with speed,
averaging about 2 fascicle lengths s-1 (at a speed of 0.64 m
s-1; see fig. 8A in Gillis and
Biewener, 2001
). In the rat, the vastus lengthened during stance
and the rate of lengthening increased with speed when walking and trotting
data are considered, averaging about 1.5-2.0 lengths s-1 (at a
speed of 0.64 m s-1; see fig. 8B in
Gillis and Biewener, 2001
).
The ankle extensors of cats have also been reported to exhibit increased
strain rates as a function of speed
(Gregor et al., 1988
;
Prilutsky et al., 1996a
;
Whiting et al., 1984
).
EMG - phase advance, duration and IEMG
EMG activity of antigravity muscles generally ends part way through stance.
In the present study of the horse, EMG activity ended at 40% of stance in the
triceps and at 59% of stance in the vastus. In the rat (see fig. 3 in
Gillis and Biewener, 2001) EMG
activity during trotting ended at about 51% of stance in the biceps femoris
and at about 63% of stance in the vastus. In the tammar wallaby, EMG activity
continued for 67% and 71% of the period of force development (plantaris and
lateral gastrocnemius, respectively) and tended to decrease with speed
(Biewener and Baudinette,
1995
).
In the triceps, EMG activity commenced an average of 0.036 s before initial
hoof contact (phase advance) and this did not change with speed. In the
vastus, EMG phase advance changed with speed from about zero seconds at low
speeds to 0.068 s at the highest trotting speed studied, averaging 0.042 s. In
the rat (Gillis and Biewener,
2001) there was a small EMG phase advance in the biceps femoris
but not the vastus lateralis, which did not change significantly with gait,
suggesting that it did not change with speed. A phase advance of about 0.050 s
was also observed in the dog semimembranosus during trotting
(Gregersen et al., 1998
) as
well as in the tammar wallaby plantaris and gastrocnemius (mean: 0.042 s;
Biewener and Baudinette,
1995
).
Muscle recruitment with speed
As predicted by hypothesis five, IEMG increased with speed in both muscles
(Fig. 4C) indicating that a
greater volume of muscle was actively recruited at higher speeds. Gillis and
Biewener (2001) reported that
EMG intensity (determined from relative average spike amplitude of the
rectified EMG signal) increased with speed in both the biceps and the vastus
of the rat. Because of the force-velocity relationship, an increase in the
volume of active muscle would be necessary if an increase in strain rate
required an increase in the number of fibers needed to produce the same force.
The increased IEMG might be explained by the speed-related increase in strain
rate of both muscles. However, an additional factor that should require an
increase in forelimb muscle force is the need to exert greater force on the
ground at faster trotting speeds. The peak vertical ground reaction force
acting on the equine forelimb increases about 23%
(Dutto et al., 2004b
) and peak
elbow moment about 60% (D. F. Hoyt, S. J. Wickler, H. M. Clayton, E. A. Cogger
and D. J. Dutto, unpublished data) as speed increases from 2.5 to 4.5 m
s-1. While the ground reaction force under the hindlimb was not
found to increase over the same range of trotting speeds
(Dutto et al., 2004b
), peak
knee moment near the middle of stance increased 18% (D. J. Dutto, D. F. Hoyt,
H. M. Clayton, E. A. Cogger and S. J. Wickler, unpublished data). Thus, the
observed increase in IEMG of both muscles appears to be consistent with
increased force development, reflecting an increased volume of active muscle
resulting from either a force-velocity effect of increased strain rate or
because of increased muscle force requirements at faster speed, or both.
Stride parameters
Duty factor differed between limbs and was longer in the forelimb than in
the hindlimb. Interestingly, at 2.5 m s-1 (a speed that is more
than 0.5 m s-1 above the walk-trot transition speed) duty factor
averaged 0.517 in the forelimb but 0.445 in the hindlimb, indicating that
there is an aerial phase between the hindlimbs but not the forelimbs. It is
well documented that bipeds run with duty factors greater than 0.5
(Gatesy, 1999;
Gatesy and Biewener, 1991
),
but it is surprising to find a quadruped trotting without an aerial phase, and
it is even more surprising that there is an aerial phase between the hindlimbs
but not the forelimbs. Two separate studies from our laboratory (Catterfeld et
al., 2002; Dutto et al.,
2004a
) have shown the same phenomenon during over-ground
locomotion.
Repeatability of sonomicrometry data
Our results show that sonomicrometry yields highly consistent data when
comparing a series of consecutive strides in the same animal and when
comparing separate studies of the same animals in different years. The
reliability (Table 5), which is
the variance ratio for ten consecutive strides in one trial averaged 0.029 and
constancy (Table 6), which is
the variance ratio over two (or more) trials, averaged 0.149. ANOVA indicated
that reliability and constancy did not vary significantly between years or
speeds. This is the first published evaluation of the constancy of
sonomicrometry data and indicates, even when data are collected from different
implants in approximately the same part of a muscle of the same animal, the
patterns are quite similar.
Conclusion
In this study we tested several hypotheses regarding the function of the
lateral triceps and vastus lateralis, two relatively long-fibered proximal
limb extensor muscles. Our data indicated that the triceps, a muscle with no
external tendon, shortened and helped cause the elbow to extend during most of
stance. The vastus, however, shortened in spite of the fact that the knee
flexed during stance, suggesting substantial stretch of its series elasticity.
The difference in the length change of the fascicles and the muscle-tendon
unit exceeded what could be accounted for by tendon strain alone, suggesting
there must be a large amount of strain in the aponeurosis through which the
muscle transmits its force to the patellar tendon. The amount of strain in the
triceps did not change with speed, but in the vastus strain increased during
both of the concentric phases of muscle activity. The strain rate and the
integrated EMG in both muscles increased with speed.
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Acknowledgments |
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References |
---|
![]() ![]() ![]() ![]() ![]() ![]() ![]() |
---|
Alexander, R. M., Maloiy, G. M. O., Hunter, B., Jayes, A. S. and Nturibi, J. (1979). Mechanical stresses in fast locomotion of buffalo (Syncerus caffer) and elephant (Loxodonta africana). J. Zool. 189,135 -144.
Back, W., Barneveld, A., van Weeren, P. R. and van den Bogert, A. J. (1993). Kinematic gait analysis in equine carpal lameness. Acta Anat. 146, 86-89.[Medline]
Back, W., Schamhardt, H. C., Savelberg, H. H. C. M., van den Bogert, A. J., Bruin, G., Hartman, W. and Barneveld, A. (1995a). How the horse moves: 1. Significance of graphical representations of equine forelimb kinematics. Equine Vet. J. 27,31 -38.[Medline]
Back, W., Schamhardt, H. C., Savelberg, H. H. C. M., van den Bogert, A. J., Bruin, G., Hartman, W. and Barneveld, A. (1995b). How the horse moves: 2. Significance of graphical representations of equine hind limb kinematics. Equine Vet. J. 27,39 -45.[Medline]
Biewener, A. A. and Baudinette, R. V. (1995). In vivo muscle force and elastic energy storage during steady-speed hopping of tammar wallabies, (Macropus eugenii). J. Exp. Biol. 198,1829 -1841.[Medline]
Biewener, A. A., Konieczynski, D. D. and Baudinette, R. V.
(1998). In vivo muscle force-length behavior during steady-speed
hopping in tammar wallabies. J. Exp. Biol.
201,1681
-1694.
Bojsen-Moller, F., Hansen, P., Aagaard, P., Kjaer, M. and Magnusson, S. P. (2003). Measuring mechanical properties of the vastus lateralis tendon-aponeurosis complex in vivo by ultrasound imaging. Scand. J. Med. Sci. Sports 13,259 -265.[CrossRef][Medline]
Carrier, D. R. (1996). Function of the
intercostal muscles in trotting dogs: ventilation or locomotion? J.
Exp. Biol. 199,1455
-1465.
Carrier, D. R., Gregersen, C. S. and Silverton, N. A.
(1998). Dynamic gearing in running dogs. J. Exp.
Biol. 201,3185
-3195.
Catterfield, G. E., Dutto, D. J., Wickler, S. J. and Hoyt, D. F. (2002). Deviant forces in trotting horses. Int. Comp. Biol. 42,1207 .
Dutto, D. J., Hoyt, D. F., Clayton, H., Cogger, E. A. and
Wickler, S. J. (2004a). Moments and power generated by the
horse (Equus caballus) hind limb during jumping. J. Exp.
Biol. 207,667
-674.
Dutto, D. J., Hoyt, D. F., Cogger, E. A. and Wickler, S. J.
(2004b). Ground reaction forces in horses trotting up an incline
and on the level over a range of speeds. J. Exp. Biol.
207,3507
-3514.
Ettema, G. J. C. (1996). Mechanical efficiency
and efficiency of storage and release of series elastic energy in skeletal
muscle during stretch-shorten cycles. J. Exp. Biol.
199,1983
-1997.
Ettema, G. J. C. (2001). Muscle efficiency: the controversial role of elasticity and mechanical energy conversion in stretch-shortening cycles. Eur. J. Appl. Physiol. 85,457 -465.[CrossRef][Medline]
Farley, C. T., Glasheen, J. and McMahon, T. A.
(1993). Running springs speed and animal size. J. Exp.
Biol. 185,71
-86.
Finni, T., Hodgson, J. A., Lai, A. M., Edgerton, V. R. and
Sinha, S. (2003). Nonuniform strain of human soleus
aponeurosis-tendon complex during submaximal voluntary contractions in vivo.
J. Appl. Physiol. 95,829
-837.
Finni, T. and Komi, P. V. (2002). Two methods for estimating tendinous tissue elongation during human movement. J. Appl. Biomech. 18,180 -188.
Fukashiro, S., Itoh, M., Ichinose, Y., Kawakami, Y. and Fukunaga, T. (1995). Ultrasonography gives directly but nonivasively elastic characteristic of human tendon in vivo. Eur. J. Appl. Physiol. 71,555 -557.
Fukunaga, T., Kawakami, Y., Kubo, K. and Kanehisa, H. (2002). Muscle and tendon interaction during human movements. Exerc. Sport Sci. Rev. 30,106 -110.[CrossRef][Medline]
Fukunaga, T., Kubo, K., Kawakami, Y., Fukashiro, S., Kanehisa, H. and Maganaris, C. N. (2001). In vivo behavior of human muscle tendon during walking. Proc. R. Soc. Lond. B 268,229 -233.[CrossRef][Medline]
Gans, C. and DeVree, F. (1987). Functional bases of fiber length and angulation in muscle. J. Morphol. 192,63 -85.[CrossRef][Medline]
Gatesy, S. M. (1999). Guinea-fowl hind-limb function. I Cineradiographic analysis and speed effects. J. Morphol. 240,115 -125.[CrossRef]
Gatesy, S. M. and Biewener, A. A. (1991). Bipedal locomotion: effects of speed, size, and limb posture in birds and humans. J. Zool. 224,127 -147.
Gillis, G. B. and Biewener, A. A. (2001).
Hindlimb muscle function in relation to speed and gait: in vivo
patterns of strain and activation in hip and knee extensor of the rat
(Rattus norvegicus). J. Exp. Biol.
204,2717
-2731.
Gregersen, C. S., Silverton, N. A. and Carrier, D. R.
(1998). External work and potential for elastic storage at the
limb joints of running dogs. J. Exp. Biol.
201,3197
-3210.
Gregor, R. J., Roy, R. R., Whiting, W.C., Hodgson, J. A. and Edgerton, V. R. (1988). Mechanical output of the cat soleus during treadmill locomotion: in vivo vs in situ characteristics. J. Biomech. 21,721 -732.[CrossRef][Medline]
Griffiths, R. I. (1991). Shortening of muscle fibers during stretch of the active cat medial gastrocnemius muscle the role of tendon compliance. J. Physiol. Lond. 436,219 -236.[Abstract]
Hershler, C. and Milner, M. (1978). An optimality criterion for processing electromyographic (EMG) signals relating to human locomotion. IEEE Trans. Biomed. Engin. 25,413 -420.[Medline]
Hof, A. L., Geelen, B. A. and van den Berg, J. W. (1983). Calf muscle moment, work and efficiency in level walking; role of series elasticity. J. Biomechanics 16,523 -537.[Medline]
Hoyt, D. F., Molinari, M., Wickler, S. J. and Cogger, E. A. (2002). Effect of trotting speed, load and incline on hind limb stance-phase kinematics. Equine Vet. J. Suppl. 34,330 -336.[Medline]
Hoyt, D. F., Wickler, S. J. and Cogger, E. A.
(2000). Time of contact and step length: The effect of limb
length, running speed, load carrying and incline. J. Exp.
Biol. 203,221
-227.
Huijing, P. A. and Ettema, G. J. C. (1988). Length-force characteristics of aponeurosis in passive muscle and during isometric and slow dynamic contractions of rat gastrocnemius-muscle. Acta Morphol. Neerlando-Scand. 26, 51-62.
Ichinose, Y., Kawakami, Y., Ito, M., Kanehisa, H. and Fukunaga,
T. (2000). In vivo estimation of contraction velocity of
human vastus lateralis muscle during `isokinetic' action. J. Appl.
Physiol. 88,851
-856.
Ito, L., Kawakami, Y., Ichinose, Y., Fukashiro, S. and Fukunaga,
T. (1998). Nonisometric behavior of fascicles during
isometric contractions of a human muscle. J. Appl.
Physiol. 85,1230
-1235.
Kawakami, Y. and Lieber, R. L. (2000). Interaction between series compliance and sarcomere kinetics determines internal sarcomere shortening during fixed-end contraction. J. Biomech. 33,1249 -1255.[CrossRef][Medline]
Kram, R. and Taylor, C. R. (1990). Energetics of running: a new perspective. Nature 346,265 -267.[CrossRef][Medline]
Maganaris, C. N. (2002). Tensile properties of in vivo human tendinous tissue. J. Biomech. 35,1019 -1027.[CrossRef][Medline]
Maganaris, C. N., Kawakami, Y. and Fukunaga, T. (2001). Changes in aponeurotic dimensions upon muscle shortening: in vivo observations in man. J. Anat. 199,449 -456.[CrossRef][Medline]
Maganaris, C. N. and Paul, J. P. (2000a). In vivo human tendinous tissue stretch upon maximum muscle force generation. J. Biomech. 33,1453 -1459.[CrossRef][Medline]
Maganaris, C. N. and Paul, J. P. (2000b).
Load-elongation characteristics of in vivo human tendon and aponeurosis.
J. Exp. Biol. 203,751
-756.
Magnusson, S. P., Hansen, P., Aagaard, P., Brond, J., Dyhre-Poulsen, P., Bojsen-Moller, J. and Kjaer, M. (2003). Differential strain patterns of the human gastrocnemius aponeurosis and free tendon, in vivo. Acta Physiol. Scand. 177,185 -195.[CrossRef][Medline]
Marsh, R. L., Ellerby, D. J., Carr, J. A., Havalee, T. H. and
Buchanan, C. I. (2004). Partitioning the energetics of
walking and running: Swinging the limbs is expensive.
Science 303,80
-83.
Monti, R. J., Roy, R. R., Zhong, H. and Edgerton, V. R.
(2003). Mechanical properties of rat soleus aponeurosis and
tendon during variable recruitment in situ. J. Exp.
Biol. 206,3437
-3445.
Muramatsu, T., Muraoka, T., Kawakami, Y. and Fukunaga, T. (2002). Superficial aponeurosis of human gastrocnemius is elongated during contraction - implications for modeling muscle-tendon unit. J. Biomech. 35,217 -223.[CrossRef][Medline]
Muramatsu, T., Muraoka, T., Takeshita, D., Kawakami, Y., Hirano,
Y. and Fukunaga, T. (2001). Mechanical properties of tendon
and aponeurosis of human gastrocnemius muscle in vivo. J. Appl.
Physiol. 90,1671
-1678.
Muraoka, T., Kawakami, Y., Tachi, M. and Fukunaga, T.
(2001). Muscle fiber and tendon length changes in the human
vastus lateralis during slow pedaling. J. Appl.
Physiol. 91,2035
-2040.
Prilutsky, B. I., Herzog, W. and Allinger, T. L.
(1996a). Mechanical power and work of cat soleus, gastrocnemius
and plantaris muscles during locomotion - possible functional-significance of
muscle design and force patterns. J. Exp. Biol.
199,801
-814.
Prilutsky, B. I., Herzog, W., Leonard, T. R. and Allinger, T. L. (1996b). Role of the muscle belly and tendon of soleus, gastrocnemius, and plantaris in mechanical energy-absorption and generation during cat locomotion. J. Biomech. 29,417 -434.[CrossRef][Medline]
Reeves, N. D. and Narici, M. V. (2003).
Behavior of human muscle fascicles during shortening and lengthening
contractions in vivo. J. Appl. Physiol.
95,1090
-1096.
Roberts, T. J. (2002). The integrated function of muscles and tendons during locomotion. Comp. Biochem. Physiol. 133A,1087 -1099.
Roberts, T. J., Marsh, R. L., Weyand, P. G. and Taylor, C.
R. (1997). Muscular force in running turkeys: The economy of
minimizing work. Science
275,1113
-1114.
Scott, S. H. and Loeb, G. E. (1995). Mechanical properties of aponeurosis and tendon of the cat soleus muscle during whole-muscle isometric contractions. J. Morphol. 224, 73-86.[Medline]
Shadwick, R. E. (1990). Elastic energy storage
in tendons: mechanical differences related to function and age. J.
Appl. Physiol. 68,1033
-1040.
Taylor, C. R. (1994). Relating mechanics and energetics during exercise. Advances Vet. Sci. Comp. Med. 38A,181 -214.
Tokuriki, M. and Aoki, O. (1995). Electromygraphic activity of the hindlimb muscles during the walk, trot and canter. Equine Vet. J. Suppl. 18,152 -155.
van Weeren, P. R., van den Bogert, A. J. and Barneveld, A. (1992). Correction models for skin displacement in equine kinematic gait analysis. Equine Vet. Sci. 12,178 -192.
Walmsley, B., Hodgson, J. A. and Burke, R. E.
(1978). Forces produced by medial gastrocnemius and soleus
muscles during locomotion in freely moving cats. J.
Neurophysiol. 41,1203
-1213.
Whiting, W. C., Gregor, R. J., Roy, R. R. and Edgerton, V. R. (1984). A technique for estimating mechanical work of individual muscles in the cat during treadmill locomotion. J. Biomech. 17,685 -694.[CrossRef][Medline]
Wickler, S. J., Hoyt, D. F., Biewener, A. A., Cogger, E. A. and De La Paz, K. L. (2005). In vivo muscle function vs speed. II. Muscle function trotting up an incline. J. Exp. Biol. 208,1191 -1200.[CrossRef]