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Footwear Affects the Behavior of Low Back Muscles When Jogging

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Abstract and Figures

Use of modified shoes and insole materials has been widely advocated to treat low back symptoms from running impacts, although considerable uncertainty remains regarding the effects of these devices on the rate of shock transmission to the spine. This study investigated the effects of shoes and insole materials on a) the rate of shock transmission to the spine, b) the temporal response of spinal musculature to impact loading, and c) the time interval between peak lumbar acceleration and peak lumbar muscle response. It was hypothesised that shoes and inserts a) decrease the rate of shock transmission, b) decrease the low back muscle response time, and c) shorten the time interval between peak lumbar acceleration and peak lumbar muscle response. Twelve healthy subjects were tested while jogging barefoot (unshod) or wearing identical athletic shoes (shod). Either no material, semi-rigid (34 Shore A), or soft (9.5 Shore A) insole material covered the force plate in the barefoot conditions and was placed as insole when running shod. Ground reaction forces, acceleration at the third lumbar level, and erector spinae myoelectric activity were recorded simultaneously. The rate of shock transmission to the spine was greater (p < 0.0003) unshod (acceleration rate: Means +/- SD 127.35 +/- 87.23 g/s) than shod (49.84 +/- 33.98 g/s). The temporal response of spinal musculature following heel strike was significantly shorter (p < 0.023) unshod (0.038 +/- 0.021 s) than shod (0.047 +/- 0.036 s). The latency between acceleration peak (maximal external force) and muscle response peak (maximal internal force) was significantly (p < 0.021) longer unshod (0.0137 +/- 0.022s) than shod (0.004 +/- 0.040 s). These results suggest that one of the benefits of running shoes and insoles is improved temporal synchronization between potentially destabilizing external forces and stabilizing internal forces around the lumbar spine.
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Int J Sports Med. 6(2001) Datei: js801 Seite: 414 25.7.2001 ± 18:05 blackcyanmagentayellow
Orthopedics and Clinical Science
Ogon M, Aleksiev AR, Spratt KF, Pope MH, Saltzman CL. Foot-
wear Affects the Behavior of Low Back Muscles When Jogging.
Int J Sports Med 2001;
22: 414± 419
after revision: October 30, 2000
Use of modified shoes and insole materials has been
widely advocated to treat low back symptoms from running
impacts, although considerable uncertainty remains regarding
the effects of these devices on
the rate of shock transmission
to the spine. This study investigated the effects of shoes and in-
sole materials on a) the rate of shock transmission to the spine,
b) the
temporal response of spinal musculature to impact load-
ing, and c) the time interval between peak lumbar acceleration
and peak lumbar
muscle response. It was hypothesised that
shoes and inserts a) decrease the rate of shock transmission,
b) decrease the low back muscle response time, and c) shorten
the time interval
between peak lumbar acceleration and peak
lumbar muscle response. Twelve healthy subjects were tested
while jogging barefoot (unshod) or wearing identical athletic
shoes (shod). Either no material, semi-rigid (34 Shore A), or
soft (9.5 Shore A) insole material covered
the force plate in
the barefoot conditions and was placed as insole when running
shod. Ground reaction forces, acceleration at the third lumbar
level, and erector spinae myoelectric activity were recorded
multaneously. The rate of shock transmission to the spine was
greater (p < 0.0003) unshod (acceleration rate: Means  SD
127.35  87.23 g/s) than shod (49.84  33.98 g/s). The temporal
response of spinal musculature following heel strike was
significantly shorter (p < 0.023) unshod
(0.038  0.021 s) than
shod (0.047  0.036 s). The latency between acceleration peak
(maximal external force) and muscle response peak (maximal
internal force) was significantly (p < 0.02
1) longer unshod
(0.0137  0.022 s) than shod (0.004  0.040 s). These results
suggest that one of the benefits of running shoes and insoles
is improved
temporal synchronization between potentially de-
stabilizing external forces and stabilizing internal forces around
the lumbar spine.
n Key words: Ru
nning injuries, shoes, low backpain, insoles.
The loading rate (the load amplitude divided by the time to
reach the maximal amplitude) has been shown to be an impor-
tant physical factor influencing injury to the musculoskeletal
system. Repetitive, rapidly applied impulsiv
e loading has been
shown to produce joint degeneration, whereas slowly applied
loads of equal or even greater magnitude often have no dele-
effects [15 ±17]. With every step, an impulsive shock
wave is generated at heel strike that is transmitted from the
lower extremities through
the spine [7]. To avoid the jarring
and potentially damaging effects of this shock wave, the hu-
man body has evolved complex mechanisms for dampening
the shock wave.
The effectiveness of shock absorbing shoes on dampening this
shock wave during gait
has been discussed controversially [3 ±
5, 7,9,11,12]. Although there is a general agreement that in-
soles or shoe modifications may lower impact loading at heel
strike [3,4, 9,12], there
is some concern that insoles may lower
the shock absorbing behavior of the body at the same time
Despite this controversy, the use of shock absorbing insoles
has been successfully used to treat
low back patients [26],
and it has been shown that shock absorbers lower the shock
wave at low back level [25].
On the other hand, an active shock absorbing behavior of the
human body has been described that might be diminished by
soft inlays [8,18, 20]. In fact, it has been shown that the devel-
opment of external forces [11], as well as the transmissibility
of impact forces through the human body [7], are increased
by wearing soft soles. So far, the mechanism of active shock ab-
sorption is unclear.
, there is some evidence that the neurophysiologic
system is involved [8,18± 21], and it has been shown that glu-
teal muscle activation in walking is related to proprioception at
the level of the sole [2]. It has also been shown that a reduced
shock absorbing capacity of the human musculoskeletal sys-
tem from the femoral condyle to the forehead correlates with
the presence of low back pain
[25]. It is interesting to note that
Footwear Affects the Behavior of Low Back Muscles When Jogging
M. Ogon
, A. R Aleksiev
, K. F Spratt
, M. H Pope
, C. L Saltzman
Department of Orthopaedic Surgery, University of Innsbruck, Austria
Iowa Spine Research Center, University of Iowa, Iowa City, USA
Department of Orthopaedic Surgery, University of Iowa Hospitals & Clinics, Iowa City, IA, USA
Biomedical Engineering Department, University of Iowa, Iowa City, IA, USA
Department of Biomedical Physics and Bioengineering, University of Aberdeen, Scottland, UK
Iowa Testing Programs, University of Iowa, Iowa City, IA, USA
Int J Sports Med 2001; 22: 414± 419
Georg Thieme Verlag Stuttgart ´ New York
ISSN 0172-4622
Int J Sports Med. 6(2001) Datei: js801 Seite: 415 25.7.2001 ± 18:05 blackcyanmagentayellow
in this study the low back patients had no X-ray findings. Thus,
there is no evidence that a damaged disc or joint is the reason
for the decreased shock
absorbing capacity, but there might be
an altered behavior of the active part of the musculosceletal
system. In fact, it has been shown that the lumbar muscles of
low back sufferers react slower to sudden loading than the
spinal muscles of healthy individuals [1]. Nevertheless, the role
of the spinal muscles in energy dissipation is as yet unclear.
With each step internal muscle forces at the low back level are
generated to achieve
equilibrium and stabilize the lumbar
spine. Proper function of this system seems crucial, because
external forces which act earlier than the human control sys-
tem is able to
respond might injure the spine. Muscle activity
dissipates energy since the contracted muscle is a viscoelastic
element. The standing impact studies of Pope et al. [14] clearly
show the attenuation
in the bent knee stance. Sudden loads
have been noted in many epiodemiologic studies to be asso-
ciated with reports of acute low back pain. If the load is applied
to a spinal motion segment, then in the absence of muscular
stability the displacement will exceed the range of the neutral
zone and soft tissue structures (e. g. interspinous ligaments
and intervertebral disc) will be loaded.
The purpose
of this study was to explore the influence of shoes
and associated materials on shock wave transmission and, si-
multaneously, the motor response in the lower back to heel
strike impact.
Based on the positive experience in low back patients with
absorbers [26] it was hypothesised that shoes would
demonstrate more protective effects than bare feet (a shoe
main effect), that soft materials would demonstrate more pro-
tective effects than hard materials or no materials (a material
main effect) and that shoes with a soft
insole would be more
protective than other combinations of shoes and materials (a
shoe by material interaction), where protective effects were
defined as:
Decreases in the loading rate experienced at the spine.
2. Decrease of the response time of the spinal muscles to heel
strike (occurrence time relative to heel strike).
3. Reduction of the time interval between peak lumbar accel-
eration and peak lumbar muscle response.
Materials and Methods
Twelve generally healthy volunteers were recruited by local
advertising. Five were female and seven male. The mean age
was 32.9 years
with a standard deviation of 7.9 years. Ages
ranged from 21 to 48. All subjects signed an informed consent
statement that had been approved by the institutions review
The experiment was a 2wShoe
3wMaterial 3wRepetition
factorial experiment where each subject ran at a self-paced
slow jogging speed on a laboratory runway of 8 m length. As a
within subjects design, each subject was exposed
to all levels of each factor in the design (jogging both with
shoes and barefoot [2wShoe], with either no material, a hard
material or a soft material placed on the for
ce plate or in the
shoe as an insole [3wMaterial] and repeating each of these six
combinations three times [3wRepetitions]). Each repetition
was considered valid when the following
criteria were met: 1)
no change in the jogging style, 2) the right heel contacted the
force plate (Type 4060A, Bertec Corporation, Worthington, OH,
USA), and
3) the velocity was constant (+/- 15 percent).
Since a fundamental assumption in repeated measures designs
the independence of events across trials, a preliminary study
was done to evaluate the independence of responses across
trials. Once the three criteria for validity of repetitions were in-
cluded, trial to trial variations in
jogging speed and mechanics
(acceleration amplitude and acceleration rate at low back lev-
el) were minimal. Further,
the low demand and simple nature
of the task made learning or fatigue effects extremely unlikely.
When subjects arrived for their scheduled participation, they
were fitted with the correct size of New Balance 600 running
shoe (New Balance Athletic Shoe, Inc. Boston, MA, USA), and
had the EMG electrodes and accelerometer devices fitted and
tested. Subjects were then informed about what was expected
of them in terms of trying to run in a natural and consistent
form across all repetitions
of the short jogging distance. Be-
cause the pilot efforts suggested little learning or fatigue ef-
fects associated with the protocol, the ordering of the trials
was fix
ed to minimize subject efforts in putting on and taking
off shoes. Subjects first nine trials were barefoot (unshod) and
their last nine trials were with shoes (shod). The nine unshod
trials varied the material on the force plate from none to hard
to soft with the material placed on the surface of the force
plate. The nine shod trials
varied the material insoles from
none to hard to soft, with the materials used as shoe insoles.
Greater detail concerning the experimental design is provided
in Table 1
Ground reaction forces, acceleration at L3 level and erector spi-
nae muscle activity were simultaneously monitored during
jogging. Ground reaction forces were measured by
a force
plate. Acceleration was recorded by a single-axis, lightweight
(0.4 g) accelerometer (Isoton, PE Accelerometer Model
2250A-10, Endevco, San Juan Capistrano, CA,
USA) attached to
the skin at the L3 spinal process with double-sided adhesive
tape [22, 27]. Low back muscle
activity was recorded 3 cm lat-
eral to the midline at L3 level by surface bipolar EMG elec-
Table 1 Experimental design
Trials Material Combination of materials with shoe
or force plate
Unshod 1± 3 None Uncovered force plate.
Unshod 4± 6 Hard Force plate covered
with 10 mm
thick, semi-rigid (35 Shore A) shock
absorbing material
Unshod 7± 9 Soft Force plate covered with 10 mm
thick, soft (9.5 Shore A) shock ab-
sorbing material.
Shod 10± 12 None Shoe (80 Shore A) without insole.
Shod 13± 15 Hard A customized, 10 mm thick insole
from semi-rigid shock absorbing ma-
terial (PE Lite
Shod 16± 18 Soft A customized,
10 mm thick insole
from soft shock absorbing material
PE Lite
, Medium Density, Knite-Ride Inc., Kansas City, MO, USA
PPT, Blue, Langer Biomechanics Group Inc., Deer Park, NY, USA
Footwear Affects Low Back Muscles when Jogging Int J Sports Med 2001; 22 415
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trodes with buildup preamplifiers to reduce the artifacts fixed
at the right side. The same location of the bipolar surface elec-
trodes had already been proven
to be the best [6,13].
Data collection for
all three systems (EMG, force plate, and ac-
readings) was triggered synchronously by a
switch, built into a ground based platform, located in the run-
way ahead of the force plate. To assess the rate of shock trans-
mission to the spine, acceleration ra
te was calculated as accel-
eration amplitude divided by the time to acceleration peak. The
time to acceler
ation peak was defined as acceleration duration
(latency between acceleration onset and acceleration peak)
(see Fig.1). To assess the low back muscle response time to heel
strike impact, the latency between heel strike and muscle re-
sponse onset (L1) was determined (see Fig. 1). To assess the la-
tency between maximal external and maximal internal force at
low back level, the
latency between acceleration peak and mus-
cle response peak (L2) was calculated (see Fig.1). The erector
spinae muscle response following heel strike was analyzed by
using inspection of digitally-magnified raw EMG signals with
resolution of 1 ms (Origin 3.5, Microcall Software Inc., North-
ampton, MA, USA), because all available time domain proces-
sing methods (average, integration, RMS, etc.) lead to an error
equal to at least the length of its time constant. Onset of muscle
response was
defined as the first increase in the EMG activity
following touchdown. Since there is also muscle activity due
to gait, only an increase of at least twice the magnitude of back
ground activity after touch down was consider
ed as muscle re-
sponse. To avoid a bias, the data were mixed in a random order
by one of the investigators and analyzed by another one.
A graphic representation of these measurements is summa-
rized in Fig.1.
A preliminary 3-way ANOVA was performed to evaluate the ef-
fects of the experimental conditions on jogging speed as an in-
itial screen for order effects. To control overall experiment-
wise error rate and to ev
aluate the overall effects of the shoe,
materials and repeated assessments on spine related out-
comes two 3-way
MANOVAs were done to simultaneously
evaluate the acceleration (amplitude and duration) and (mus-
cle onset latency [L1] and latency from acceleration peak to
muscle maximum res
ponse [L2]) criteria. A critical p-value of
0.10 was used to identify significance for the overall MANOVA
procedures. Follow-up univariate 3-way ANOVAs for the indi-
vidual outcomes, and
for the acceleration rate criteria were
planned for any of the significant MANOVA results, and post
hoc multiple comparisons were performed using Tukeys high-
est significant difference (HSD) follow-up procedures. Signifi-
cant interactions among the experimental factors in the ANO-
VA procedures were followed up using tests of simple effects
under the assumption that all of the factors in the
models re-
presented fixed effects. Critical p-values of 0.05 were used to
identify significant results from follow-up
univariate ANOVAs
and follow-up tests of main and simple effects.
Regarding the shock transmission to the spine, the overall 3-
way MANOVA demonstrated
significant shoe and material ef-
fects by Wilks Lambda, F
= 57.03, p < 0.0001 and F
= 6.69,
p < 0.0003, respectively. Based on the 3-way ANOVA follow-
ups, acceleration amplitude demonstrated no significant shoe
(p = 0.18), material (p = 0.35), or repetition (p = 0.42) main ef-
fects, nor any interaction effects involving these factors. How-
ever, acceleration duration demonstrat
ed a significant shoe
main effect (p < 0.0001), with shorter durations in the unshod
(0.021 0.0
09 s) compared to the shod conditions (0.039
0.011 s). A significant material main effect (p < 0.0006), with
shorter durations in the no material condition (0.027
0.014 s) compared to
the hard and soft material conditions
(0.030 0.013 s and 0.033 0.014 s, respectively) was found.
The materials themselves were not significantly different from
each other based on Tuke
ys HSD follow-ups of the significant
materials main effect. The pattern of results for acceleration
rate were consistent with those observed for acceleration
duration. There was a significant shoe main effect (p < 0.0003),
where acceleration
rate was significantly higher unshod
(127.35 87.23 g/s) compared to shod (49.84 33.98 g/s). The
significant materials effect for acceleration rate (p < 0.02)
again demonstrated a similar
grouping pattern. Significantly
higher deceleration rates in the no material conditions
(106.79 91.42 g/s) compared to
the hard and soft material
conditions (78.07 61.56 g/s and 80.27 71.09 g/s, respective-
ly) were found, which were not themselves significantly differ-
ent from each other based on Tukeys HSD
follow-ups of the
significant materials main effect. Acceleration rate measured
in the different material conditions, barefoot and shod, are
presented in Fig. 2.
Fig.1 Accelero-
meter and EMG sig-
nal from one
recorded simulta-
neously with the
force plate meas-
urement. Heel strike
was determined by
the force plate.
Acc = acceleration,
Resp = erector spine
muscle response,
= acceleration
duration, L1 = laten-
cy between heel
strike and muscle
response onset,
L2 = latency be-
tween acceleration
peak and muscle re-
sponse peak.
Fig. 2 Acceleration
rate at low back lev-
el following heel
strike in shod and
unshod conditions.
No material (no),
hard shock absorb-
ing material (hard),
and soft
shock ab-
sorbing material
(soft) was placed
on the ground and
in shoes, respective-
ly. Standard error
bars are indicated.
Int J Sports Med 2001; 22 Ogon M et al416
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The muscle onset latency (L1) and the latency between accel-
eration peak and muscle peak response (L2) were considered
simultaneously within a 3-way MANOVA, which
neously considered the effects of shoe, materials, repetitions
and their interactions. Results revealed a significant shoe main
effect by Wilks Lambda, F
= 10.81, p < 0.0034. The materials
and repetition main effects were non-significant, nor were any
of the interaction effects significant.
Delay in the temporal response of the spine musculature to
strike (L1) showed a significant shoe main effect
(p < 0.023.) The low back muscle response onset following heel
strike occurred significantly earlier in the unshod (0.038
0.021 s)
than shod condition (0.047 0.036 s). There was a
trend toward a material main effect (p < 0.064) with the pat-
tern of means in the expected order with fastest onset of mus-
cle response with no material (0.036 0.021 s), a slower re-
sponse with hard material (0.041 0.020 s), and the slowest
response with soft material (0.050 0.042 s) as shown in Fig. 3.
The latency between peak acceleration and peak muscle re-
sponse at the lower back showed a significant shoe main effect
(p < 0.021) with the pattern of
means demonstrating signifi-
cantly shorter latencies in the shod conditions (0.0040
0.040 s) compared with the unshod
conditions (0.0137
0.022 s) (Fig. 4). No significant material (p = 0.62) or repetition
(p = 0.39) main effects or interaction effects were observed.
The average jogging velocity was 1.49 m/s (range, 1.45 m/s to
1.51 m/s). Jogging velocity was not significantly related to
shoe, material, or
This study was conducted to address the hypotheses that
wearing shoes and insert materials 1) decreases the rate of
shock transmission to the lower back, 2) decreases the re-
sponse time of the spinal muscles to
heel strike, and 3) reduces
the time interval between acceleration peak and muscle re-
sponse peak at the
lower back in jogging. The results support
hypothesis 1 and 3. Hypothesis 2 was not supported. On the
contrary, the muscle response was significantly later in run-
ning shod than in running barefoot and, furthermore, in-
creased with increasing softness of the sole material.
The study hypotheses were based on the positive experience
with shock absorbing materials in low back patients
[15]. From
this report it was assumed that shock absorbers shorten the la-
tency between external (passive shock, potentially destruc-
and internal (active muscular, protective) forces experi-
enced at the low back. The observed neuromuscular delay
caused by shoes and insert materials suggests, on the first
view, that this assumption was wrong. However, at the same
time, the latency between heel strike and acceleration peak at
low back level increased by wearing shoes, due
to an increased
time interval between acceleration onset and acceleration
peak. Thus, with shoes, the lower back experienced the impact
force peak later.
In fact, this mechanical delay predominates
the muscle response delay so that, after all, wearing shoes de-
creases the time interval between maximum external and
maximum internal forces
experienced in the lower back dur-
ing running. Since the shock wave tends to destabilize equilib-
rium and the muscle response may help regain equilibrium, it
seems that shortening of this time interv
al may have clear
The results raised the question whether the proprioception af-
ferent from the heel, or low back muscle stretch reflex triggers
the neuromuscular shock-absorbing behavior at lumbar level.
The muscle responses occurred with a mean latency between
36 and 50 ms after touchdown. This is long enough for the
stretch reflex (M1), which takes 30 to 50 ms
[8, 23]. The onset of acceleration at the low back occurred
about 18 ms after touchdown, and the latency between accel-
eration onset and muscle response onset was under 20 ms in
the barefoot situation, which is too
short even for an M1 reflex.
Based on these temporal data it appears that the lumbar mus-
cle activity was not triggered by a local shock wave at the lum-
bar level, but by proprioception
at the heel level. The proprio-
ceptors appear sensitive enough to discriminate differences of
the magnitude of the standard
deviation of the impact dura-
tion (about 0.01 s). This indicates that heel afferent proprio-
ception could be sensitive to heel loading rate.
Perhaps the single biggest limitation of this study is the lack of
multiple shoe conditions, which kept this study from consider-
the possible differential effects of various shoe designs rel-
ative to various introductions of insole materials. The inability
Fig. 3 Latency between heel strike and erector spinae muscle re-
sponse onset (L1). Resp Onset = muscle response onset, Resp End =
muscle response end. Standard error bars are indicated.
Fig. 4 Latency between acceleration peak (Acc Peak) and muscle re-
sponse peak (Resp Peak) following heel strike in shod and unshod
condition (L2). Standard error bars are indicated.
Footwear Affects Low Back Muscles when Jogging Int J Sports Med 2001; 22 417
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of the current study to demonstrate the anticipated interaction
between the shoe (shod, unshod) and material (none, hard,
soft) factors was, on first consideration, a surprise.
The most
obvious explanation, lack of power owing to the relatively
small sample size of twelve subjects, although plausable and
almost obligatory,
was unsatisfying as the pattern of results
across the material conditions in the shod and unshod condi-
tions did not hint toward any anticipated effects. Similarly,
confounding effects due to lack of randomization of the order-
ing of the shoe/material combinations was not judged a likely
explanation for either the observed shoe and mat
erial main ef-
fects or for the lack of the anticipated shoe x material interac-
tion effects. The jogging task was simply too similarly executed
across trials and
shoe/material combinations to suggest either
learning or fatigue effects.
Another limitation is the fact that it is extremely difficult to
measure the force that acts in vivo on the lumbar spine direct-
ly. However, accelerometers have been proven useful to esti-
mate impact loading. T
o reduce the bias in the acceleration
outcomes, we followed the fundamental requirements to
properly measure bone vibration in vivo by skin attached ac-
ers: a thin layer of soft tissue between accelerometer
and bone [22], use of a light weight accelerometer [10], and a
tight attachment to the skin [22,2
Even if there is latency difference in the other
mutually ortho-
gonal directions at the lumbar level, it would
not significantly
influence the results and conclusions of the study, because of
the too brief latency of the axial component of the accelera-
tion. The few significant differences in external
forces between
the hard and the soft insole materials should also be interpret-
ed with some caution. In this study we investigated two of the
most commonly used inshoe orthotic materials - PeLite and
PPT. These materials are US orthotic industry
standards, and
represent the range of softness-hardness that can be used
comfortably without adversely affecting gait. It might be ex-
pected that testing harder and softer insole mate
rial than
these we used might provide a more robust test of the hypoth-
eses. In this initial study
, however, scientifically well charac-
terized materials were used for the sake of the clinical signifi-
cance, practical application and future reproduction of the
study design.
Robbins and Gouw [19], in a review of the litera-
ture and their own results on athletic footwear and chronic
overloading, summarized that soft shoes with thick yielding
midsoles are probably dangerous because they
attenuate tac-
tile plantar sensations required for protective impact moderat-
ing behavior. They developed the theory of a plantar surface,
sensory-mediated feedback system
for neuromuscular control
of the shock absorbing behavior [18,20, 21]. Their theory has
been supported by the experimental observation that stereo-
typic ipsilateral
hip flexion and contralateral hip extension fol-
lowing a rapid, heavy loading of the leg increases in amplitude
as the irregularity of the plantar surface support increases
The low back muscle response to running was dramatically af-
fected by use of shoes or soft insole shoeing materials. The la-
tency from heel
strike to muscle response onset was pro-
longed. This is consistent with the observation of Robbins and
Gouw who found that the active damping of the external force
energy started later and progressed less efficiently with
use of
shoes or soft insole materials [18,20]. Thus, too soft shoes
might be worse, especially for low back patients since their
ability to react to sudden load
is diminished anyway [1].
Our results help to explain other observations of shock absorb-
ing behavior [7, 24]. Forner et al. [7] recently examined the
properties of shoe insert materials as they
effect shock wave
transmission between tibia and forehead. They studied the dif-
ference between materials with lower rigidity and loss tangent
(low energy absorbing) and higher rigidity with
high loss tan-
gent. They found more transmission of acceleration from the
tibia to the forehead with the least rigid material. Our study
suggests that this decrease in shock absorbing behavior
is due
to an increased latency of spinal muscle response when wear-
ing very soft shoes.
In summary, we found that shoes and insert materials not only
reduce the loading rate, but affect the low
back muscles. They
can protect the lower spine from heel strike impact in two
ways: by reducing the impact loading rate and by minimizing
latency between maximum external and internal force.
This study was supported by an Erwin-Schroedinger Grand
(Austria), NIH (G 50111), and the University of Iowa (P10542).
The authors
thank Donald Shurr, PT, C.P.O., for his expert ortho-
tic advice. Presented at the 53. AAOS (American Academy of
Orthopaedic Surgeons) Meeting. Specialty Day of the
(American Orthopaedic Foot and Ankle Society). San Francisco,
CA, USA, February 16th, 1997.
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Corresponding Author:
Associate Professor Michael Ogon, M. D.
Department of Orthopaedic Surgery
University of Innsbruck
Anichstrasse 35
A-6020 Innsbruck
Phone: +43 (512) 504-2697
Fax: +43 (521) 50
Footwear Affects Low Back Muscles when Jogging Int J Sports Med 2001; 22 419
... Impact-related shocks can be attenuated passively via soft tissues, footwear and ground substrate compliance, or actively via muscles that do negative work or modify gait kinematics (Paul et al., 1978;McMahon et al., 1987;Derrick et al., 1998;Whittle, 1999;Edwards et al., 2012;Butler et al., 2003;Addison and Lieberman, 2015). For example, several studies have explored how footwear and foot strike patterns affect impact shock transmission through the musculoskeletal system (Ogon et al., 2001;Divert et al., 2005;Lieberman et al., 2010;Kulmala et al., 2013;Boyer et al., 2014;Gruber et al., 2014;Giandolini et al., 2016). However, few studies have examined impact-related shocks in the spinal column. ...
... lower limb and head), and these studies have mostly focused on timedomain rather than frequency-domain accelerations (e.g. Voloshin and Wosk, 1982;Ogon et al., 2001;Delgado et al., 2013). ...
... Study participants were barefoot during experiments in order to reduce SA from passive structures such as shoes (Paul et al., 1978;Whittle, 1999;Addison and Lieberman, 2015). Participants were instructed to use a rearfoot strike pattern during running trials to control for potential variation in lumbar shock transmission and temporal response of spinal musculature caused by different foot strike patterns (see Ogon et al., 2001;Delgado et al., 2013). Once instrumented with accelerometers and motion-tracking markers (see below), participants stood motionless in a neutral position with their arms resting comfortably at their sides for 30 s while 3D kinematics measured static standing posture. ...
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During locomotion, each step generates a shock wave that travels through the body toward the head. Without mechanisms for attenuation, repeated shocks can lead to pathology. Shock attenuation (SA) in the lower limb has been well studied, but little is known about how posture affects SA in the spine. To test the hypothesis that lumbar lordosis (LL) contributes to SA, 27 adults (14 male, 13 female) walked and ran on a treadmill. Two lightweight, tri-axial accelerometers were affixed to the skin overlying T12/L1 and L5/S1. Sagittal plane accelerations were analyzed using power spectral density analysis, and lumbar SA was assessed within the impact-related frequency range. 3-D kinematics quantified dynamic and resting LL. To examine the effects of intervertebral discs on spinal SA, supine MRI scans were used to measure disc morphology. Results showed no associations between LL and SA during walking, but LL correlated with SA during running (p<0.01, R2=0.30) resulting in as much as 64% reduction in shock signal power among individuals with the highest LL. Patterns of lumbar spinal motion partially explain differences in SA: larger amplitudes of LL angular displacement and slower angular displacement velocity during running were associated with greater lumbar SA (p=0.008, R2=0.41). Intervertebral discs were associated with greater SA during running (p=0.02, R2=0.22), but after controlling for disc thickness, LL remained strongly associated with SA (p=0.001, R2=0.44). These findings support the hypothesis that LL plays an important role in attenuating impact shocks transmitted through the human spine during high-impact, dynamic activities such as running.
... During human locomotion, contacts of the feet with the ground produce impact forces that cause a shock wave to travel through the musculoskeletal system from foot to head (Lafortune et al., 1996;Ogon et al., 2001). It has been speculated that one of the purposes of running footwear is to reduce these impact shocks (Hagen and Hennig, 2009;van Mechelen, 1992). ...
... It has been speculated that one of the purposes of running footwear is to reduce these impact shocks (Hagen and Hennig, 2009;van Mechelen, 1992). However, the effectiveness of shock-absorbing shoes on reducing this shock wave during gait remains controversial (Ogon et al., 2001). ...
Running is a modality that has a large number of adepts, including women. Therefore, it is important to understand how sportswear can help women, with special attention to the breast movement. The aim of this study was to analyse the effect of different combinations of breast support and footwear on the breast movement during walking and running. Twenty women performed treadmill walking (5 km/h) and running (7 and 10 km/h) combining different footwear (barefoot, minimalist, and traditional) and breast support conditions (bare breast, everyday bra, and sports bra). Three-dimensional data from breast and trunk markers were tracked for 10 stride cycles. Relative breast displacement was calculated and derived for velocity. An interaction effect was observed between support, footwear, and speed conditions. The bare breast conditions presented differences from the other conditions in the majority of the kinematic variables, presenting higher breast displacement and velocity values. On the other hand, the sports bra conditions presented the lowest values for the kinematics variables. In the vertical component of breast displacement during running (10 km/h) we verified that the sports bra presented reductions of 56% and 43% in relation to the bare breast and everyday bra conditions, respectively. Despite this, no differences were found between footwear within each breast support condition. A sports bra is efficient to decrease breast movement. In addition, neither of the tested footwear was able to decrease these movements.
... However, minimalist footwear or even barefoot running may reflect a lower injury risk and a more economical running pattern (Lieberman, 2012;Tam et al., 2014). The effectiveness of shock-absorbing shoes on reducing this shock wave during gait has been controversial (Ogon et al., 2001), analysing the differences between different models and their relationships with lower limb injuries, with little attention to their effects on spine posture. ...
... In the Bfoot condition, the lumbar region was less rectified than Trad or Min footwear conditions, suggesting that both footwear conditions were able to maintain the natural curvatures of the spine on the lower back. These findings corroborate previous studies that demonstrated that both traditional and minimalist footwear are able to reduce the impacts of running (Bonacci et al., 2013;Lieberman, 2012;Ogon et al., 2001). However, future research should investigate if changes in spine kinematics are related to the capability of the impact reduction of both footwears. ...
... Five sets of 20 drop jumps were performed with a 10-s interval between jumps and a 2-min rest period between sets. Drop jumps were performed barefoot, because it has been reported that biomechanical factors are influenced dramatically by shoes (De Wit et al., 2000;Ogon et al., 2001) and because eliminating the effect of shoes was thought to facilitate comparisons between sets. ...
... This load can cause micro-trauma to the underlying tissues and may eventually cause permanent damage to the legs. Besides the bottom part of the body, the shock will also transmit to upper part, resulting in maladies of the ankle, knee, hip, and even the lower back [3][4][5][6]. Some sorts of injuries can be permanent, especially at the fragile knee joints part. ...
... Furthermore, the use of shoes and insole materials could cause the shock transmission to the spine if the material are not suitable 13 . This study stated their concerned about the insole that may have low shock absorbing behaviour. ...
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Working in prolonged standing position among industrial workers has been shown to be associated with different potentially serious health outcomes, namely lower back pain, leg pain, fatigue, discomfort, and other health issues. Personalisation of insole offers a solution that will provide a perfect fit and comfort to the shoes wearer based on the ergonomic considerations. It works in a way that it alters the pressure away from painful areas by increasing the surface area that supports the weight of the body and evenly distributes it to the whole plantar area. Survey was conducted among workers at a manufacturing industry company to study on the level of pain experienced by them together with their foot anthropometry. Then, the foot pressure of each of the workers was collected by using pressure measurement device (F-scan). Combination of these data was used to design the customized insole that is fit for the worker. The personalised insoles were fabricated by using Additive Manufacturing technology. After that, the insoles were validated by using the F-scan and Electromyogram (EMG) to ensure their effectiveness in reducing pressures on the foot and muscle activity hence improving the comfort of the shoe wearer. At the end of the experiment, it was found that the insole is able to reduce the peak pressure of four out of five areas of the worker’s foot with the reduction of pressure percentage ranging from 6% to 28%.
Objective: Amputation of the lower limb due to loss of part of the musculoskeletal structure reduces performance and increases injury during locomotion. The effect of various types of prosthetic feet has been analyzed in several studies during running. The purpose of this study was a biomechanical analysis of the influence of SACH and Dynamic-Response foot on several kinetic variables in the stance phase of running in individuals with unilateral transtibial amputation. Materials & Methods: In this semi-experimental study, 8 left foot transtibial amputees were included in this study using an available or easy sampling method. The target population was unilateral transtibial amputees who were able to run and the available population included left transtibial amputees who were referred to Kosar Rehabilitation Center in Tehran from 2008 to 2012. To adapt to the foot, each foot was used by the subjects for at least one week before the experiment. All subjects participated in 3 running evaluation sessions; 1 session involving the use of their own foot (familiarization session), 1 session involving the use of SACH foot, and 1 session involving the use of Dynamic foot. Only data from the 2 last sessions were used to compare both feet. Each subject runs in 12-meter walkway 3 times at a speed of 2.5 meters per second. The same running speed was chosen for the comparability of kinetic variables. Sport shoes were used to bring the test conditions closer to the actual running conditions. In each session, 3 successful trials were performed so that the foot was in full and perfect contact with the force plate. Kistler force plate and three-dimensional motion analysis Vicon system were used to collect kinetic and kinematic data, respectively. The motion and the force plate data were sampled simultaneously at 200 and 1000 Hz, respectively. The trajectories of the markers and analog data were filtered using the predicted mean square error adaptive filter in version 1.7 of the Vicon software package. The Kinetic variables were generated using the dynamic model of the Vicon Plug-in-Gait. The vertical ground reaction force was normalized for body weight. In the present study, 5 variables were selected for biomechanical analysis of feet. The maximum vertical ground reaction force, power, spring efficiency, ankle moments at the amputated leg, and the symmetry ratio (percentage) of the maximum vertical ground reaction force between the amputated leg and the intact leg were calculated. All values in each trial were averaged for each subject with each foot. A paired t-test and a Wilcoxon test were used to analyze the data based on normality (P ≤0.05). Results: In examining the normality of the data distribution, the results showed that the data of maximum power absorption of the ankle with the SACH foot and the maximum power absorption of the hip with the Dynamic-Response foot did not have a normal distribution and other variables had a normal distribution. The results of paired t-test and Wilcoxon showed that Spring Efficiency and Maximum Plantar Flexion were significantly different between the SACH and Dynamic-Response feet (P ≤0.05). The Spring Efficiency was greater with Dynamic-Response foot than the SACH foot (P =0.05), although the Maximum Plantar Flexion with the SACH foot was greater than Dynamic-Response foot (P =0.05). While there is no statistical difference between the maximum vertical ground reaction force, maximum power absorption and generation in the ankle, maximum power absorption and generation in the knee, maximum power absorption and generation in the hip, maximum dorsiflexion moment, and the symmetry ratio (percentage) of the maximum vertical ground reaction force between the amputated leg and the intact leg. Conclusion: The results of the study showed that the spring efficiency with Dynamic-Response foot was greater than SACH foot and closing to the spring efficiency of a normal foot. With this perspective, the Dynamic-Response foot has more natural performance than the SACH foot.
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Wearable technologies are often indicated as tools that can enable the in-field collection of quantitative biomechanical data, unobtrusively, for extended periods of time, and with few spatial limitations. Despite many claims about their potential for impact in the area of injury prevention and management, there seems to be little attention to grounding this potential in biomechanical research linking quantities from wearables to musculoskeletal injuries, and to assessing the readiness of these biomechanical approaches for being implemented in real practice. We performed a systematic scoping review to characterise and critically analyse the state of the art of research using wearable technologies to study musculoskeletal injuries in sport from a biomechanical perspective. A total of 4952 articles were retrieved from the Web of Science, Scopus, and PubMed databases; 165 were included. Multiple study features—such as research design, scope, experimental settings, and applied context—were summarised and assessed. We also proposed an injury-research readiness classification tool to gauge the maturity of biomechanical approaches using wearables. Five main conclusions emerged from this review, which we used as a springboard to propose guidelines and good practices for future research and dissemination in the field.
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Inertial measurement units (IMUs) can be used to monitor running biomechanics in real-world settings, but IMUs are often used within a laboratory. The purpose of this scoping review was to describe how IMUs are used to record running biomechanics in both laboratory and real-world conditions. We included peer-reviewed journal articles that used IMUs to assess gait quality during running. We extracted data on running conditions (indoor/outdoor, surface, speed, and distance), device type and location, metrics, participants, and purpose and study design. A total of 231 studies were included. Most (72%) studies were conducted indoors; and in 67% of all studies, the analyzed distance was only one step or stride or <200 m. The most common device type and location combination was a triaxial accelerometer on the shank (18% of device and location combinations). The most common analyzed metric was vertical/axial magnitude, which was reported in 64% of all studies. Most studies (56%) included recreational runners. For the past 20 years, studies using IMUs to record running biomechanics have mainly been conducted indoors, on a treadmill, at prescribed speeds, and over small distances. We suggest that future studies should move out of the lab to less controlled and more real-world environments.
Background and significance: A controversy persists about the role of pelvic unlevelness and leg length inequality (LLI) as etiologic and aggravating factors in low back pain (LBP), and the diagnostic approach to the use of heel lifts. A question arises: does LLI causes LBP, or is pelvic unlevelness a result of LBP? If the latter, why would we try to change the posture by unilateral heel lift, presumably to something more painful in LBP patients? Purpose: The aims of this study was: a) to investigate the external forces by means of force plate analysis, and the internal forces of the spine by erector spine surface EMG during isometric contraction and sudden load, and b) to define how these responses were modified with or without expectancy and before and after fatigue, when artificial LLI was created in normals and in chronic LBP patients. Subjects and methods: In the first study, 10 patients with chronic LBP (age 41.4, SD 9.6) and 10 matched healthy subjects (age 41.1, SD 9.4) were investigated. The patients participated in a well established 2-week rehabilitation program. The pain degree was quantified by Visual Analogue Scale (VAS). The subjects stood on a force platform with extended knees, their arms along the body and pelvis against a board to push off. In this position they extended their back against a mechanical resistance at 20% of the individual maximum voluntary contraction MVC for 4 s. Surface EMG activity of the multifidus and longissimus were recorded. An artificial LLI was created by placing different boards of 12 mm and 25 mm respectively under the left and right foot in random order. After the 2-week rehabilitation course the same investigation was repeated for the LBP patients. The same procedures were repeated for the controls after 2 weeks. In the second study, 11 chronic LBP patients (7 males 38.4 years SD 9.8, and 4 females 37.2 years SD 3.6) and 11 age and sex matched controls (7 males with mean age 39.5 SD 9.8, and 4 females with mean age 36.2 SD 3.7) were recruited. The experimental setup was the same as in the first study. Expected and unexpected load and unloads were applied before and after fatigue at the level of T4 by weight of 2 kg attached via a load cell to a harness around the subject's shoulder. The weight was dropped from a height of 45 cm, applying a sudden forward bending moment. Results: The results from the first study showed that in healthy subjects EMG activity increased proportionally on the side with a board and decreased correspondingly on the contralateral side. In LBP patients the EMG increment is higher when the artificially elongated leg corresponds to the more painful side. In healthy subjects COP is close to the middle line, and shifts proportionally away from the longer leg side. In LBP patients COP is shifted initially away from the more painful side even without artificial LLI. COP shifts significantly more in patients when the longer leg corresponds to the more painful side. The displacement of COP is significantly smaller at the end of the rehabilitation program for chronic LBP. The healthy subjects did not show electrophysiologic signs of muscle fatigue, detected by median frequency shift, after 45 s of isometric contraction at 20% MVC. The LBP patients not only were fatigued under the same conditions, but showed asymmetric erector spine fatigue, higher on the side corresponding to the longer leg. The fatigue index decreased significantly after the rehabilitation program. The most fascinating result in the second study was a 5-level interaction of LId together with expectation, fatigue, group, and axis, defined from the force plate parameters. The results of the EMG reaction time and magnitude to sudden load were expressed mostly by a 3-level interaction of LLI together with expectation and group. An obvious discrepancy between the ground reaction forces (expressing the external forces) and the EMG activity (showing the internal forces) comparing normals versus LBP patients were found. The EMG magnitude to sudden load was smaller but the magnitude of the ground reaction forces were larger in LBP patients versus normals. EMG reaction time was slower but the latency of the ground reaction forces are faster in LBP patients versus normals. This internal/external force discrepancy increased after placing a heel lift under the foot, corresponding to the painful side in LBP patients. Conclusion: This study proved only the short term beneficial effect of this frontal plane postural correction in chronic LBP patients. Further studies are necessary to verify a longer term effect of monitored frontal plane posture correction in chronic LBP.
The effect of the soft tissue between bone and a preloaded skin surface accelerometer was studied in vivo by comparing its output with the output of an accelerometer connected directly to the bone by a needle through the soft tissue. A 34-g skin surface accelerometer gave an output with little resemblance to the bone motions, appearing to oscillate at its resonant frequency on the soft tissue. A 1. 5-g skin surface accelerometer showed nearly identical output to the bone acceleration.
A method for measuring the transmissibility of the human spinal column to vertical vibrations using light-weight, skin-mounted accelerometers is described. The accelerometers were conveniently attached to the skin at the S2 and T2 levels of the spine using adhesive tape. The acceleration time records were analysed using a discrete Fourier transform to calculate the amplitude for each frequency component up to 40 Hz. Transmissibility was estimated as the ratio of the output over the input for each frequency component. The analysis included a compensation for both skin movement and differences in the inclination of the spine from the vertical at the accelerometer mounting sites. The healthy spine was observed to attenuate frequency components above 20 Hz, whereas in ankylosing spondylitis the spine behaved as a rigid strut.
The feasibility of a pendulum impact method to establish the dynamic response of the standing subject was explored. Threaded k-wires were placed in the L3 lumbar spinous process and in the posterior superior iliac spine. The gain and phase angle between the platform and the vertebra were established. The lower extremities were found to be very important in the attenuation of the impulse, while different shoes had little ability to attenuate the resonance peak.
Analytical investigations of impact absorption of linear and isothermal viscoelastic materials are described. Three methods based on different considerations and approximations are studied, and similarities are shown in their results. For a viscoelastic buffer of a given thickness, the optimal loss tangent is determined to be approximately one. Greater reductions in impact force can be achieved if the high loss is accompanied by stiffness reduced by a factor of three to four compared with that of an elastic buffer. If the impactor is spherical rather than flat, a higher loss tangent, of the order of 10, is needed to minimize the impact force. A more sophisticated interpretation scheme for the ball rebound test for screening the loss tangent of viscoelastic materials is derived.
Studies of the relation between joint function and mechanical stress have led to a revival of the old concept that primary osteoarthritis is actually a wearing out of joints. Recent experimental evidence suggests that joints can wear out by repetitive impulsive loading, rather than by rubbing. This new mechanistic approach is compatible with the pathology of, and clinical experience with, the disease.
The vibration response of long bones to the impact of an instrumented hammer was monitored with an accelerometer. Different parameters associated with the accelerometer waveform were evaluated with respect to their sensitivity to changes in the thickness and quality of simulated soft tissue over excised bones and to changes in the preload force between the accelerometer and bone. Soft-tissue related errors were reduced if the accelerometer was spring-loaded rather than mass loaded. Experiments with human volunteers indicated no saturation of soft-tissue related effects at any preload level below the pain threshold.
This study focused on the electromyographic activity of the trunk musculature, given the well-documented link between occupational twisting and the increased incidence of low back pain. Ten men and 15 women volunteered for this study, in which several aspects of muscle activity were examined. The first aspect assessed the myoelectric relationships during isometric exertions. There was great variability in this relationship between muscles and between subjects. Further, the myoelectric activity levels (normalized to maximal electrical activity) obtained from nontwist activities were not maximal despite maximal efforts to generate axial torque (e.g., rectus abdominis, maximum voluntary contraction; 22% external oblique, 52%; internal oblique, 55%; latissimus dorsi, 74%; upper erector spinae [T9], 61%; lower erector spinae [L3], 33%). In the second aspect of the study, muscle activity was examined during dynamic axial twist trials conducted at a velocity of 30 and 60 degrees/s. The latissimus dorsi and external oblique appeared to be strongly involved in the generation of axial torque throughout the twist range and activity in the upper erector spinae displayed a strong link with axial torque and direction of twist, even though they have no mechanical potential to contribute axial torque, suggesting a stabilization role. The third aspect of the study was comprised of the formulation of a model consisting of a three-dimensional pelvis, rib cage, and lumbar vertebrae and driven from kinematic measures of axial twist and muscle electromyograms. The relatively low levels of normalized myoelectric activity during maximal twisting efforts coupled with large levels of agonist-antagonist cocontraction caused the model to severely underpredict measured torques (e.g., 14 Nm predicted for 91 Nm measured). Such dominant coactivity suggests that stabilization of the joints during twisting is far more important to the lumbar spine than production of large levels of axial torque.