Vol. 90, Issue 1, 205-215, January 2001
Static and dynamic postural control in long-term microgravity:
evidence of a dual adaptation
Guido
Baroni1,2,
Alessandra
Pedrocchi1,2,
Giancarlo
Ferrigno1,2,
Jean
Massion3, and
Antonio
Pedotti1,2
1 Centro di Bioingegneria, Politecnico di Milano, Fondazione
Don Carlo Gnocchi, Istituto di Ricovero e Cura a Carattere
Scientifico, I-20148 Milan; and 2 Dipartimento di
Bioingegneria, Politecnico di Milano, I-20133 Milan, Italy; and
3 Laboratory of Neurobiology and Movements, Centre National
de la Recherche Scientifique, 13402 Cédex 20 Marseille, France
 |
ABSTRACT |
The adaptation of dynamic
movement-posture coordination during forward trunk bending was
investigated in long-term weightlessness. Three-dimensional
movement analysis was carried out in two astronauts during a 4-mo
microgravity exposure. The principal component analysis was applied to
joint-angle kinematics for the assessment of angular synergies. The
anteroposterior center of mass (CM) displacement accompanying trunk
flexion was also quantified. The results reveal that subjects kept
typically terrestrial strategies of movement-posture coordination. The
temporary disruption of joint-angular synergies observed at subjects'
first in-flight session was promptly recovered when repetitive sessions
in flight were analyzed. The CM anteroposterior shift was consistently
<3-4 cm, suggesting that subjects could dynamically control the
CM position throughout the whole flight. This is in contrast to the
observed profound microgravity-induced disruption of the quasi-static
body orientation and initial CM positioning. Although this study was
based on only two subjects, evidence is provided that static and
dynamic postural control might be under two separate mechanisms,
adapting with their specific time course to the constraints of microgravity.
motor control; posture; sensorimotor adaptation; motion analysis
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INTRODUCTION |
IT IS WELL KNOWN THAT
exposure to weightlessness causes major changes in the
sensory-perceived environment and in external constraints, such as
those related to static and dynamic equilibrium control. As a
consequence, the quantitative description of the adaptation of the
human motor system to the sustained gravitational-altered condition
represents a unique opportunity for better identification of the
variables primarily controlled in postural control and detailed
description of the complex role of gravity in movement-posture coordination (15).
When the experimental observation extends for a long period of time,
such as for several months of weightlessness exposure, the complete
process of sensorimotor adaptation to the unusual environment can be
described (19). Long-term analysis of human motor behavior
in weightlessness may provide particularly enlightening information for
motor rehabilitation, where functional recovery depends on the
patients' ability to learn new motor strategies that are compatible
with their permanent lesions. It is also crucial for future space
missions when astronauts will be required to work in weightless
conditions and perform demanding skillful tasks over long periods of time.
In a recent study, the static control of posture was quantitatively
investigated in two subjects during 4 mo of microgravity exposure for
the first time (3). Results confirmed the strong effect of weightlessness on the control of whole body static
orientation already obtained in the frame of short-term space missions
(4, 22, 23) and parabolic flights (17, 30).
However, the long-term observation revealed that the anteroposterior
(AP) position of the body center of mass (CM), which is a fundamental
reference for postural control in normogravity (21), was
involved in a long-term process of adaptation throughout the entire
flight toward the reemergence of typical ground-based CM positioning
compatible with equilibrium. Although this result was obtained on a
limited data set in flight, it suggested that a long-term process of
adaptation to the weightless environment might lead to the reemergence
of normogravity strategies of quasi-static postural control based on CM
AP regulation.
In light of this result, the present study addresses the question of
the long-term adaptation of the dynamic control of posture in
weightlessness. The aim was to check whether gradual adaptation processes to microgravity, which are comparable to those observed for static posture, were present, suggesting that static and dynamic postural control might be ruled out by the same mechanism. In contrast,
different modes of long-term adaptation would indicate that static and
dynamic postural regulations are governed separately, reacting
specifically to the prolonged microgravity exposure.
Related investigations of the dynamic movement-posture coordination
during short-term space missions and parabolic flights pointed out a
possible different adaptation of static and dynamic postural regulation
to weightlessness. Clement et al. (4) reported that the
unaltered postural adjustments accompanying arm raising and standing on
tiptoe performed by two subjects during the few days of their
permanence in orbit were observed, despite a marked disruption of the
initial static whole body position. In a later study, during a 7-day
spaceflight, Clement et al. (5) reported forward-biased,
quasi-static postural regulation, together with preserved anticipatory
and compensatory postural adjustments. Further experimental evidence
was provided by Massion et al. (22, 23) and
Vernazza-Martin et al. (35), who observed unaltered movement coordination and dynamic CM position control, despite a
persistent, biased trunk orientation and variable initial CM positioning during short-term space missions and microgravity exposure
on parabolic flights.
Although only two subjects took part in the reported analysis of
long-term microgravity, our results provide further evidence that
terrestrial movement-posture coordination strategies based on CM
position control are typically unchanged during the whole duration of
spaceflight. This occurs despite the observed bias affecting the
quasi-static regulation of whole body orientation (3) and
the temporary impairment of the kinematic synergies observed for both
subjects at their first inflight session.
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METHODS |
Experimental design.
Two subjects, one aged 40 yr (subject A) and the other 38 yr
(subject B), took part in the study. Subject A
was on his second spaceflight, whereas subject B was on his
first mission in orbit. The experiments were performed on board the
core module of the Russian space station Mir, during the 179-day
European Space Agency mission Euromir95.
The test subject was firmly fixed to the floor of the space module with
Velcro shoes and straps and was asked to stand in the upright position
with his hands clasped behind his back. On the command "go," he was
instructed to bend his trunk forward by 30° along the anatomic AP
direction and then return to the initial position. The movement had to
be performed as fast as possible. The forward trunk bending was
followed by a symmetrical backward trunk inclination, which this paper
does not deal with. The movement was performed both with eyes open (EO)
and closed (EC). Ten trials were acquired for each experimental condition.
For subject A, inflight data were collected on flight
day 11 (FD11), FD19, FD69, and FD113; for
subject B, experimental sessions were performed only late in
flight on FD150. An on-ground baseline data set was acquired for both
subjects 17 days before flight (F
17). Postflight sessions were
performed 1 day after reentry (R+1) for both subjects, on R+3 for
subject B and on R+5 for subject A. Only data
acquired in these latter postflight sessions were used as the
postflight data sets in this study.
Data collection and kinematic analysis were carried out by using a
space-qualified version of the ELITE automatic opto-electronic motion
analyzer (13), working with passive markers at a 50-Hz sample rate (12). Seven retroreflective markers
were applied to the subjects' skin using biocompatible adhesives.
Markers were positioned on easily identifiable anatomic landmarks
and/or bony processes, as shown in Fig.
1. The marker repositioning error between
different experimental sessions was estimated to be <1.5 cm
(3).
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Fig. 1.
Marker arrangement model and corresponding stick diagram
used for upper trunk forward-bending quantitative description. Joints
considered for angle analysis are reported ( , ankle;  , knee;  ,
hip;  , trunk). Markers were placed on the following anthropometric
landmarks: temple (1), zygoma (2), acromion
(3), anterior-superior iliac spine (4),
greater trochanter (5), lateral femoral condyle
(6), and lateral malleolus (7).
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The two-dimensional marker coordinates that were measured on the ground
and during flight underwent on-ground postprocessing consisting of
tracking, three-dimensional (3D) reconstruction (13), and
filtering (7). System accuracy was assessed as equal to
1/1,700 of the working volume diagonal (3,285 mm), meaning a error in
locating the 3D positions of the markers of <2 mm (12).
Angle analysis.
Angle analysis was carried out on the ankle, knee, and hip joints. For
each trial, the movement start was defined as follows
where Xi is the position of the marker on
the acromion (see Fig. 1). Movement onset was, therefore, indicated
when the difference between the current (n + 1) position of
the marker and the mean value of the previous n positions
became equal or greater than twice the standard deviation (95% of the
population) on the previous n frames. The end of the
movement was found in a similar manner, taking into account the phase
in which the subject recovered the erect posture before the execution
of the backward-leaning movement.
Joint angles were analyzed on the anatomic sagittal plane, which was
reconstructed from the 3D movement description (3). The
anatomic AP axis (
) was analytically defined as follows
where the symbol × denotes cross product,
is the vertical direction, and 
is the normal of the
regression plane among the positions of the markers placed on the
acromion and great trochanter (see Fig. 1) during forward bending.
As indicated in Fig. 1, the ankle angle () was taken as the angle
between the line connecting the knee to the lateral malleolus and the
AP axis; the knee angle () was taken as the angle between the line
connecting the femoral condyle to the greater trochanter and the line
connecting the femoral condyle to the lateral malleolus; and, according
to Massion et al. (22, 23), the hip angle () was taken
as the angle between the line connecting the acromion to the iliac
spine (trunk axis) and the line connecting the greater trochanter to
the femoral condyle. Trunk inclination () was also measured and was
taken as the angle between the trunk axis and the vertical axis. The
range of angular motion (ROM) was measured as the difference between
the joint angles at the initial position and at the maximum forward
trunk bending.
Principal component analysis.
Principal component (PC) analysis was applied on ankle (), knee
(), and hip () joint angles (see Fig. 1) to assess their linear
correlation over time (1, 14, 20). PC analysis transforms the vector that represents the time variation of the three joint angles
(t), (t), (t) around their
average values (µ
,
µ
,µ
) into a sum of orthogonal time-dependent components
[PCi(t), where i = 1, 2, 3] weighted
with three constant orthonormal vectors with components
wi,j (where i,j =
1, 2, 3). Analytically, this transformation is given by
![<FENCE><AR><R><C>[&agr;(t)−&mgr;<SUB>&agr;</SUB>]</C></R><R><C>[&bgr;(t)−&mgr;<SUB>&bgr;</SUB>]</C></R><R><C>[&ggr;(t)−&mgr;<SUB>&ggr;</SUB>]</C></R></AR></FENCE>](/content/vol90/issue1/fulltext/205/img006.gif)
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(1)
|
where T denotes the transpose operation, and the
vectors wj
[wj = (w1,j
w2,j w3,j)T]
are the PC loadings that correspond to the eigenvectors of the jth eigenvalue of the correlation matrix of the data.
PCi (i = 1, 2, 3) gives the minimum number of
statistically independent linear components that describe the
time-dependent processes. Therefore, if trunk bending is performed
according to certain intrinsic kinematic constraints on the joint
angles, then the number of PCs sufficient to describe the movement with good approximation will be less (possibly even only 1) than the number
of kinematic degrees of freedom (3 in this case). In our analysis, we
used the percentage of total angular variance (TAV) described by the
first PC as a way of measuring how accurately the movement could be
described using only one linearly independent variable, namely the
first PC (PC1). We called this index the PC1 factor and obtained it
analytically using the following formula
![PC1 factor<IT>=</IT>[var (<IT>w<SUB>11</SUB> ∗ </IT>PC<IT>1</IT>)<IT> + </IT>var (<IT>w<SUB>21</SUB> ∗ </IT>PC<IT>1</IT>)<IT>+</IT>var (<IT>w<SUB>31</SUB> ∗ </IT>PC<IT>1</IT>)]<IT>/</IT>TAV<IT> ∗ 100</IT>](/content/vol90/issue1/fulltext/205/img009.gif)
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(2)
|
where var is variance, and TAV is defined as the sum of the
angle variances over time. In addition, we used the PC loadings as a
way to quantify the level of encoding of each angular time process in
the specific PC. Indeed, according to Eq. 1,
vector wj = (wi,j
wi,j
wi,j)T accounts
for the contribution of the jth PC to the specific
joint-angular variation (i = 1, 2, 3).
It is important to stress here that, unlike Mah et al.
(20), who explicitly assumed to include the effects of
signal standard deviations in their PC analysis, we carried out a
preliminary normalization by dividing the time course of each angle by
its standard deviation. This was done to mask out the effect of
different signal amplitudes. The relevance of this effect is clear if
we consider that the total angular variation described by PC1 (PC1 factor) on three random processes with standard deviations comparable to trunk-bending joint angles was found to be 88%. After
normalization, this value was reduced to the expected 33%.
CM kinematics.
The displacement of the CM with respect to the ankle-joint axis was
estimated by using a seven-segment biomechanical model. The position of
the CM and the mass of the various body segments were taken from
anthropometric tables (8, 36). The model was validated
according to the protocol proposed by Rabuffetti and Baroni
(31) during various standing activities performed by a
pool of control subjects (axial movements, leg raising, squatting, and
jumping on the spot). The model was also tested in quasi-static conditions by comparing the model-estimated CM projection with the
position of the center of pressure (CP), which was measured by using a
piezoelectric force platform. The results showed high correspondence
along the AP axis (CPAP CMAP = 0.20 ± 1.64 mm). In addition, the theoretical CM shift, which
would have been caused only by the forward trunk bending in absence of
any postural adjustment at the lower limb level, was estimated
according to Vernazza-Martin et al. (35) and was used to
evaluate the effects of the kinematic synergies on the actual CM AP
shift. For this aim, the following CM compensation index was calculated
as (35)
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where CMtheoretical is the model-predicted CM shift
that would have occurred if no postural corrections had been produced, and CMactual is the model-predicted CM shift accounting for
the measured postural adjustments.
Statistical analysis.
A one-way between-group variance analysis (ANOVA) was carried out with
the aid of the software package Statistica (StatSoft, Tulsa, OK). The
hypotheses of the ANOVA model were checked by assessing the data
fit to the normal distribution (Kolmogorov-Smirnov and 
2
tests) and the homogeneity of the variances (Levene's test). The level
of significance was confirmed each time by using the nonparametric
between-group ANOVA-equivalent Kruskal-Wallis ANOVA by ranks. Specific
effects were evaluated by using Scheffé's post hoc comparisons
of means. The null hypothesis was rejected when probability fell below
0.05.
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RESULTS |
Movement kinematics.
Figures 2 and
3 report, for subjects A
and B, respectively, movement kinematics (exemplifying stick
diagrams and joint-angle time courses), the theoretical AP CM
displacements (the one that would have been caused by trunk bending
only, see METHODS), and the actual CM AP shift with respect
to the initial position. The corresponding joint ROM at the peak
forward-bending position (average ± SD) is shown in Table
1.
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Fig. 2.
Stick diagram of upper trunk-bending trials that show the
on-ground and in-flight movement-posture coordination strategies of
subject A. Dotted sticks, initial posture; solid sticks,
peak forward position. Stick trajectories have been represented only
for the markers involved in the analysis (upper trunk, lower trunk,
thighs, and shanks). The joint-angle time courses and the resulting
center of mass (CM) anteroposterior slope measured with respect to the
initial position are reported. For this representation, each trial was
normalized in time, taking as reference value the average movement
duration of the specific experimental session. The actual CM shift is
compared with the theoretical CM displacement, which the trunk
displacement would have caused in absence of postural countermovements.
F17, 17 days before flight; FD11, FD19, FD69, FD113: flight
days 11, 19, 69, 113; R+5,
5 days after reentry.
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Fig. 3.
Stick diagram of upper trunk-bending trials that show the on-ground
and in-flight movement-posture coordination strategies of subject
B. Dotted sticks, initial posture; solid sticks, peak forward
position. Stick trajectories have been represented only for the markers
involved in the analysis (upper trunk, lower trunk, thighs, and
shanks). The superimposition of joint-angle time courses for most
of the considered trials at each session is reported. The resulting CM
anteroposterior shift measured with respect to the initial position is
compared with the theoretical CM displacement. A procedure of
intrasession time normalization of the reported trials (see Fig. 2) was
performed.
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The stick diagram representation highlights the marked backward bias of
the initial posture for subject A on FD11 and subject B on FD150. According to Baroni et al. (3), the
progressive recovery of the initial erect posture disruption is evident
for subject A, giving rise by FD113 to an in-flight initial
posture compatible with equilibrium requirements. This observation is not confirmed for subject B, who still exhibited, late in
flight, a backward-biased initial posture accompanied by a forward
inclination of the trunk. Despite the evident bias in static postural
regulation with marked differences between the two subjects, it is
evident that, during the dynamic phase of the movement, both subjects displaced the lower body segments (thighs and pelvis) in the opposite direction with respect to the forward trunk bending at all sessions in
flight. Thus the typically terrestrial synergistic production of the
prime movement and postural adjustments turned out to be consistently
used throughout the whole mission. These unaltered "axial
synergies" efficiently compensated the dynamic CM displacement caused
by the trunk bending, as qualitatively represented by the theoretical
and actual CM shift time courses.
For subject A, the amplitude of trunk inclination (angle in Fig. 1) corresponding to the achievement of the required task was
significantly reduced early on in the flight (FD11) but was gradually
regained during the mission. Statistical analysis of angle between
different experimental sessions (ANOVA with post hoc comparison)
revealed a gradually decreasing significance in the differences between
preflight and in-flight sessions. On FD113, the ROM of angle was
found not to differ significantly (P = 0.82) with
respect to the reference on-ground data. The trunk ROM was found to be
highly sensitive to the presence of visual cues in microgravity. Both
subjects showed significantly larger trunk-bending amplitudes with
their eyes shut during all in-flight sessions [ANOVA with visual
condition as independent variable, lowest significance
F(1,17) = 5.52, P = 0.031 for subject A on FD19], except on FD113, when a
significantly larger range of trunk motion was exhibited by
subject A (P = 0.03) in the EO condition.
Kinematic synergies.
The evaluation of the postural adjustments accompanying trunk bending
revealed that the plantarflexion of the ankle-joint angle ()
remained close to preflight values and was quite independent of the
initial body configuration, which varied during the flight (see Fig. 2
and cf. Fig. 3). Statistical analysis on angle ROM confirmed the
lack of significant differences between the normogravity and
microgravity ankle motion for both subjects [subject A:
F(1,34) = 1.12, P = 0.29;
subject B: F(1,14) = 2.67, P = 0.12].
Noteworthy is the different knee-joint kinematics between the two
subjects in flight. Whereas subject A still extended his knee when bending forward, subject B flexed his knee on
reaching the peak forward position. Indeed, the specificity of the
knee-joint kinematics was found to be the major factor influencing the
level of movement coordination, which was significantly decreased at the subjects' first in-flight session, as reported in Fig.
4.
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Fig. 4.
Principal component (PC) analysis results. PC1 factor, i.e.,
the percentage of total angle variance expressed by the first principal
component, is reported for both subjects. * Statistical significance
(P < 0.05) compared with normogravity values (F17).
Subject A regained terrestrial preflight values by FD19 and
maintained them throughout the whole 4-mo flight and on return to
Earth. Subject B still performed movements in a poorly
coordinated way late in flight on FD150 during his first in-flight
session.
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Before flight, the strong coupling in the time of joint-angular
variation shown by subject A (mean PC1 factor is 98.9 ± 0.8%) is opposed to a slight, lower coordination with higher
variability exhibited by subject B (PC1 factor = 93.9 ± 5.2%). When trunk bending was performed in microgravity,
both subjects showed a significantly reduced level of coordination
among joint angles with respect to the on-ground reference data, as
indicated by the considerable decrease in the PC1 factor at their first
inflight session (PC1 factor = 83.4 ± 8.6% for
subject A on FD11 and 69.1 ± 10.7% for subject B
on FD150). Starting from his second experimental session on FD19,
subject A exhibited a prompt recovery of the joint-angular
synergies, which were maintained for the rest of the flight.
Statistical analysis of the PC1 factor for subject A (post
hoc comparison with test session as independent variable) revealed that
the overall statistical significance
[F(4,23) = 13.7, P = 0.000007] was mainly attributable to significant differences between
preflight and early in-flight (FD11) sessions (P = 0.000023).
A decreased movement coordination was also confirmed by statistics for
subject B, when the PC1 factor was compared between preflight and in-flight trials [F(1,10) = 13.1, P = 0.004]. At the postflight sessions, both
subjects highly coordinated the postural adjustments with the prime
movement, showing PC1 factor values close to 100%, with no significant
differences with respect to preflight trials [subject A:
F(1,7) = 0.7, P = 0.4;
subject B: F(1,4) = 1.07, P = 0.4].
The comparative representation of exemplifying joint angles and PC1,
PC2, and PC3 time courses (Fig. 5)
suggests that, in the presence of high coordination (on F17 and on
FD113 for subject A; preflight for subject B)
(Fig. 5, left; see Fig. 4), the kinematic synergy among hip,
knee, and ankle joints is accounted for by the single component PC1,
with PC2 only related to the initial slight knee flexion (on F17 for
subject B) (35) and with PC3 at noise level. By
contrast, when movement-posture coordination was low (on FD11 for
subject A; on FD150 for subject B) (Fig. 5,
right; see Fig. 4), its decrease is mainly attributable to the knee-joint kinematics, as indicated by the similar time courses of
knee-angular variation and PC2. In this case, movement execution is
governed by two degrees of freedom: PC1, accounting for hip and
ankle-joint covariation, and PC2, related to the additional independent
factor represented by the knee-angular variation.
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Fig. 5.
Comparison of hip, knee, and ankle time courses and PC
time courses for exemplifying trials at all sessions. For the trials
with high-movement coordination (see PC1 factor preflight for both
subjects and on FD113 for subject A), the overall
joint-angular variation is well interpreted by the first principal
component (PC1), being PC2 and PC3 at noise level. The decrease in
movement coordination (see PC1 factor on FD11 and FD150) is explained
by the independent knee-joint kinematics, which is well described by
the second principal component (PC2). The motor task is here
characterized by 2 kinematic degrees of freedom.
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This result is generalized in Fig. 6, in
which PC1 loadings for hip and knee angles are reported. As introduced
in METHODS, PC1 loadings quantify the contribution of PC1
to the specific joint-angular variation, thus representing the level of
encoding of each joint kinematics accounted by PC1. For both subjects, the PC1 loading associated with the hip angle (PC1 loading) is always negative and close to unity, thus accounting for the main part
of hip angular variation at all sessions. The negative sign is due to
the increasing time of the PC1 time course, as opposed to a time
decrease of angle (see Fig. 5).
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Fig. 6.
PC1 loadings for hip (PC1 loading) and knee (PC1 loading).
The emergence of a second degree of freedom for the total angular
variation description is associated with a decrease (on FD11 for
subject A) or a change in sign (on FD150 for subject
B) of the PC1 loading related to the knee joint (PC1 loading).
Increased PC1 loadings on FD19 and FD69 denote a stronger and highly
coordinated contribution of the knee-joint kinematics to the overall
movement execution.
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Important changes involved the PC1 loading associated with the
knee angle (PC1 loading) for both subjects at their first in-flight
session. Subject A showed significant reduction of PC1 loading on FD11, denoting decreased coordination of the knee joint with
the prime movement. The same result was found for subject B
on FD150, when the change in sign accounted for an in-flight switch
from the on-ground knee extension (pre- and postflight) to the knee
flexion accompanying the forward bending (see Figs. 3 and 5).
Noteworthy is the increase in PC1 loading shown by subject
A on FD19 and FD69. The reason is that the more flexed initial
posture at knee level characterizing these two in-flight sessions (see
Fig. 2 and cf. Fig. 3) implies the performance of a more pronounced and
highly coordinated knee extension with respect to the trials on ground
and on FD113, which are characterized by a more extended initial
postural attitude (see Fig. 2 and cf. Fig. 3).
CM kinematics.
The CM kinematics associated with trunk bending are summarized in Fig.
7.
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Fig. 7.
Top: anteroposterior positioning of the CM with respect
to the ankle-joint axis (absolute position) at the initial posture
( and dashed line) and at the maximum forward leaning
( and solid line). Bottom: anteroposterior
CM displacement at the peak forward bending measured with respect to
the CM initial position (relative shift). The size of the CM relative
shift associated with trunk flexion was <4 cm for both subjects.
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Interestingly, despite the large backward bias in absolute CM
positioning during the flight (see Fig. 7, top, and cf. Fig. 3), the final CM position at the peak forward leaning was found to be
displaced very little with respect to its initial position (see Fig. 7,
bottom). The amount of compensation and/or overcompensation of the CM theoretical shift, which would have been caused by the trunk
forward bending in the absence of postural corrections (see Figs. 2 and
3), is related to the ratio of the range of trunk flexion to the size
of the postural adjustments. The corresponding values of the CM
compensation index are reported in Table
2.
The in-flight mean CM compensation index accounts for a marked
overcompensation of the CM shift. Indeed, both subjects showed a
dynamic backward displacement of the CM, which is opposite in direction
with respect to the theoretical forward CM shift. This caused a
markedly higher CM compensation index (>100%) than the one measured
before flight and in postflight sessions (always <100%) (see Figs. 2
and 3). It is worth noting that the CM overcompensation observed in
flight does not imply higher postural destabilization. Indeed, values
of the CM compensation index between 100 and 200% account for a lower
CM displacement (in the opposite direction) with respect to the CM
theoretical shift.
Despite the change in direction with respect to on-ground trials, the
CM relative shift in microgravity exhibited a stable pattern, with
absolute values ranging between 3 and 4 cm with respect to the initial
position. This is a striking result if one considers the significant
changes between on-ground and in-flight sessions affecting the initial
joint-angle configuration (see Figs. 2 and 3), the level of movement
coordination (see Fig. 4), and the relative contribution of each joint
angle to the whole movement production (see Figs. 5 and 6). When
statistics were performed on the absolute values of the CM shift in
normogravity and microgravity, no significant differences were found
for subject A [F(5,25) = 2.030403, P = 0.1]. The same result was found when the
outcomes were compared between the two subjects
[F(4,29)=2.342850, P = 0.08]. It is as if the overall movement organization, whose features appear to be subject specific and depend on the time course of
sensorimotor adaptation during flight, would guarantee, in any case, to
keep the CM shift associated with the trunk bending within the narrow
limits of dynamic whole body balance.
In addition, when the AP relative CM displacement was statistically
analyzed in EO and EC conditions, no significant differences were
found. This result suggests a rather independent CM control from the
presence of visual cues, which is in contrast to the observed
sensitivity to vision found for the prime movement kinematics.
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DISCUSSION |
The first consideration is that the reported analysis is based on
a data set from only two subjects. This is a common problem when
experiments involving human subjects are performed in space. This
limitation hinders us from generalizing our results and drawing definitive conclusions on the adaptation of human postural control to
microgravity. On the other hand, this investigation provides a unique
experimental contribution toward the deeper understanding of human
motor behavior in prolonged weightlessness and, possibly, represents a
driving factor for future related experimental activities in the frame
of the forthcoming International Space Station exploitation phase.
From this perspective, it is worth noting that consistent results
between the two investigated subjects were obtained when the
performances of their respective first experimental sessions in-flight
were compared (FD11 for subject A; FD150 for subject B). Both subjects were found to keep the CM AP position close to
that of the initial erect posture (see Figs. 2, 3, and 7), suggesting
that they were able to maintain efficient control of the CM AP
positioning during the whole flight duration. In addition, both
subjects consistently exhibited typically terrestrial postural adjustments, despite relevant changes affecting the kinematics of the
prime movement (trunk ROM and velocity; see Figs. 2 and 3). At their
first experimental sessions in flight, both subjects showed markedly
modified joint-angular synergies (PC1 factor; see Fig. 4) and the
emergence of a second degree of freedom, mainly related to the
knee-joint kinematics (see Figs. 5 and 6). For subject A,
who performed repetitive sessions in flight, the coupling among joint
activation was found to be recovered, starting from his second
in-flight session on FD19, suggesting that kinematic synergies might be
rapidly regained as an efficient strategy for multijoint movement
execution in microgravity.
The performance of forward trunk bending appeared to be markedly
slower. Early in flight (subject A on FD11), this was due to
the decreased movement amplitude associated with a rough preservation of the movement duration. A reduced ROM is a common observation in
microgravity (18) and is interpreted as related to a
reduced sensitivity of the spindle afferents and a reduced gain of the myotatic loop, resulting in lower agonist muscle activation. Late in
the mission (subject A on FD113; subject B on
FD150), the long-term sensorimotor adaptation allowed subjects to
recover the required trunk range of motion; however, this was
accompanied by a longer time required for the movement execution (see
Table 1, Figs. 2 and 3). The slowing down of the movement in
weightlessness might depend on the need to shift from a predominantly
feed-forward mode of control to a predominantly feedback mode, mainly
useful for the fine tuning of the CM trajectory during the trunk movement.
Regardless of the modifications of trunk flexion kinematics along the
flight, both subjects consistently exhibited typically terrestrial
postural adjustments involving ankles, knees, and pelvis at all
sessions (see Figs. 2 and 3). This resulted in the efficient
compensation of the CM displacement, which remained confined within a
few centimeters from the beginning to the end of the flight (see Fig.
7) and was irrespective of the CM positioning at the initial erect
posture, which deeply changed throughout the flight (see Fig. 7 and cf.
Fig. 2). This result is in agreement with the outcomes of previous
investigations during parabolic flights and short-term space missions,
in which microgravity was reported to deeply affect the static body
orientation and initial CM positioning (inducing subject-specific
forward or backward biases) (3) but to have little
influence on dynamic movement-posture coordination and CM position
control (4, 5, 22, 23, 35).
The different sensitivity of static and dynamic control of
posture and CM positioning might originate from the involvement of
different sensory cues for quasi-static whole body orientation and for
the dynamic movement-posture coordination. On one hand, the somesthetic
perception, with respect to the external environment, mainly relies on
integrating the information received from vision and various types of
gravity-related sensory cues [otoliths and truncal graviceptors
(24); muscle force proprioceptors (9); cutaneous sole plantar inputs (16)] and on
microgravity-biased muscle spindle length and velocity-sensing spindle
afferents (18). On the other hand, dynamic sensory cues,
such as those sensing the mass distribution and inertial properties of
body segments (34), the labyrinthine information
(2), and the dynamic cutaneous (16) and
proprioceptive (33) inputs, are unaffected by exposure to
microgravity and are used to achieve proper kinesthetic perception in a
weightless environment. When this unbiased sensory support is counted
on, centrally encoded, stereotyped strategies of CM displacement
control would be consistently used in flight for dynamic
movement-posture coordination during trunk movement. The dynamic CM
control would be organized, at least partly, in a closed-loop mode with
respect to the initial CM AP position, which would continue to serve as
an egocentric frame of reference for motor control (11),
despite its static bias with respect to the external environment (see
Figs. 2, 3, and 7).
This view was questioned by Pozzo et al. (30), who
doubted the role of CM positioning for postural control on the basis of
findings in short-term microgravity. They reported that subjects asked
to perform whole body lifting tasks exhibited an initial backward-biased CM positioning followed by a markedly increased dynamic
CM shift with respect to normogravity. One interpretation of these
contrasting results is that the sensorimotor adaptation to
weightlessness might be task dependent. The goal-directed task represented by whole body lifting could be performed on the basis of
different reference values with respect to the trunk-bending task. An
alternative and not exclusive interpretation is that the complexity of
the whole body lifting task would require a longer time for adaptation,
comparable with the long-term adaptation described for static CM
positioning, which only recovered the ground-based value at the end of
the flight (3).
According to our results on upper trunk bending, the only noticeable
difference between normo- and microgravity CM dynamic control is the
direction of the residual CM shift: for both subjects, the CM moved
forward in normogravity and backward in microgravity. We do not feel
confident to speculate on this result, apart from noting that the CM
relative displacement in microgravity was consistently confined within
narrow limits. The observed change in direction could be interpreted as
an overcompensation of the CM shift due to the reduced amplitude of the
trunk movement and the preservation of the postural adjustments (see
Table 2) (35). This could also depend on the
backward-leaning initial posture (see Figs. 2 and 3), as well as the
observed increased role of the knee joint in task execution. However,
backward shift is not specific to microgravity. In normogravity, it was
shown that the direction of the CM shift was subject dependent,
observing both forward and backward CM shifts (1). In
addition, when similar experiments were performed on parabolic flights,
only two of the five recorded subjects showed a backward CM shift when
trunk forward bending was performed during the 0-g phase of
the parabola (35).
Interestingly, subject A accomplished CM displacement
control in weightlessness by regaining and progressively refining
typically terrestrial motor strategies based on synergistic joint-angle activation. Indeed, strong kinematic synergies were described as
underlying the minimization of the CM shift during trunk bending in
normogravity (6, 21, 26-28), as well as during
short-term microgravity exposure [parabolic flights (35),
short-term space missions (22, 23)]. The characteristic
coupling among hip-, knee-, and ankle-joint movements, illustrated by
the PC analysis in normogravity (1) and on parabolic
flights (35), was, in this case, deeply modified during
the first in-flight recording session for both subjects, regardless of
the time elapsed since their exposure to weightlessness (see Fig. 4).
The decrease in movement coordination was characterized by a strong
representation of a second component (PC2) mainly related to the knee
movement (see Fig. 5). This result could be interpreted as indicating
that, in the presence of a biased internal representation of the body configuration, due to the impaired sensory cues in a microgravity environment and the low familiarity with the motor task, the subjects adjust the CM position continuously by using movements of the knee
joint, which present the lowest inertial characteristic compared with
the ankle and hip joints (25), comparable to a skier
during a fast descent. The impaired kinematic synergy underwent
short-term adaptation, which was most probably favored by the
increasing familiarity with the specific protocol. As early as at the
second in-flight recording (for subject A), the strong
coupling between the angle changes (PC factor >95%; see Fig. 4) was
recovered and served to refine the overall movement organization toward
a more efficient task production, which is in agreement with the
provided instruction.
The last consideration concerns the possible origin and purpose of the
observed postural adjustments in weightless conditions, where there is
no need for equilibrium control. Let us first consider the theory that
the unaltered opposite displacements of lower body segments that
accompany voluntary trunk movements and the resulting compensation of
CM shift are simply an effect induced by passive dynamic interactions
between segments when trunk bending is performed. As shown by the
modeling study of Ramos and Stark (32), these interactions
in normogravity would have provoked backward falling in the absence of
the anticipatory tibialis anterior activation seen at the onset of the
trunk bending (6). On the basis of experimental
observation in normogravity and modeling studies, it was concluded that
the postural adjustments cannot have a purely passive origin but must
be produced by centrally controlled commands that are mainly focused on
ankle flexor muscles. Comparable conclusions were drawn by Eng et al.
(10) when they modeled the dynamic interactions between
segments during arm raising. Although no comparable modeling studies on
the interactions between segments during trunk bending were performed
in microgravity, it would be surprising that, in the very different
gravitational context and in the presence of the observed modification
of prime movement kinematics, the kinematic synergies would result
solely from the passive effect of the dynamic interactions between
segments (1).
In conclusion, the reported analysis in long-term microgravity,
although based on the analysis of only two subjects, provides evidence
that the static control of the CM AP position during erect posture and
the dynamic control of the CM position during trunk bending might
depend on two different control mechanisms. The quasi-static CM
positioning appears to be deeply disrupted by microgravity exposure and
eventually exhibits a long-term adaptation toward the restoring of the
ground-based AP CM position compatible with equilibrium
(3). By contrast, the dynamic CM control during trunk
bending is maintained during the whole spaceflight. As on Earth, the
dynamic movement posture coordination produces an evident minimization
of the CM shift, regardless of subject-specific biases affecting the
initial whole body configuration. The idea that the observed dynamic CM
stabilization mainly serves as a way of counteracting the inertial
perturbing effects of prime movement performance is a fascinating topic
for further investigations. The role of postural adjustments in
dynamically regulating body stability may indeed be maximized in
microgravity, where the need to maintain equilibrium is no longer a
constraint. This idea is in accordance with evidence that motor
coordination aims at minimizing the displacement of the CM projection
(and, with good approximation, also the CP). Indeed, this would reduce
the variation of the "arm" of the total moment, both of reactive
and inertial forces around the ankle joint. From this point of view,
the main function of postural strategies in space could be interpreted
as serving to reduce movement-induced perturbations at the interface
between the human body and the external environment, with implications regarding the energy required to produce the movements themselves (29).
| |
FOOTNOTES |
Address for reprint requests and other correspondence: G. Baroni, Centro di Bioingegneria, Via Capecelatro, 66, I-20148 Milano, Italy (E-mail: baroni{at}biomed.polimi.it).
The costs of publication of this
article were defrayed in part by the
payment of page charges. The article
must therefore be hereby marked
"advertisement"
in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.
Received 1 August 2000; accepted in final form 9 August 2000.
| |
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ABSTRACT
INTRODUCTION
