Estimate Muscle Forces Research Articles

Effect of stretching training on the viscoelastic properties of human tendon structures in vivo.

The purpose of this study was to examine whether stretching training altered the viscoelastic properties of human tendon structures in vivo. Eight men performed the stretching training for 3 wk. Before and after the stretching training, the elongation of the tendon and aponeurosis of medial gastrocnemius muscle was directly measured by ultrasonography while the subjects performed ramp isometric plantar flexion up to the voluntary maximum, followed by a ramp relaxation. The relationship between the estimated muscle force (Fm) and tendon elongation (L) during the ascending phase was fitted to a linear regression, the slope of which was defined as stiffness of tendon structures. The percentage of the area within the Fm-L loop to the area beneath the curve during ascending phase was calculated as an index representing hysteresis. To assess the flexibility, the passive torque of the plantar flexor muscles was measured during the passive stretch from 0 degrees (anatomic position) to 25 degrees of dorsiflexion with a constant velocity of 5 degrees/s. The slope of the linear portion of the passive torque-angle curve during stretching was defined as flexibility index. Flexibility index decreased significantly after stretching training (-13.4 +/- 4.6%). On the other hand, the stretching training produced no significant change in stiffness but significantly decreased hysteresis from 19.9 +/- 11.7 to 12.5 +/- 9.5%. The present results suggested that stretching training affected the viscosity of tendon structures but not the elasticity.

Measurement of viscoelastic properties of tendon structures in vivo

The purpose of this study was to investigate the viscoelastic properties of tendon structures in humans. Elongation of the tendon and aponeurosis of medial gastrocnemius muscle (MG) was directly measured by ultrasonography, while subjects (n=19) performed ramp isometric plantar flexion up to the voluntary maximum, followed by a ramp relaxation. The relationship between tendon elongation (L) and estimated muscle force (Fm) was fitted to a linear regression, the slope of which was defined as compliance of the tendon structures. The hysteresis was calculated as the ratio of the area within the L-Fm loop (elastic energy dissipated) to the area beneath the load portion of the curve (elastic energy input). The resulting L-Fm relationship was non-linear in form, as previously reported on animal and human tendons in vitro. The mean compliance was 4.5+/-1.1. 10-2 mm/N. However, there was a considerable inter-subject variability (2.9 to 7.2. 10-2 mm/N). The Young's modulus, i.e., the slope of the stress-strain curve, was 280 MPa, which tended to be lower than the previously reported values for human tendons. It was also found that the strain of the tendon structures was homogeneously distributed along their length. The mean hysteresis (energy dissipation) was 22.2+/-8.8%. However, again there was a considerable inter-subject variability (9.7 to 37.2%). The present results indicated that the tendon structures of human MG were considerably compliant and their hysteresis was in accordance with previously reported values.

Effects of resistance and stretching training programmes on the viscoelastic properties of human tendon structures in vivo.

The present study examined whether resistance and stretching training programmes altered the viscoelastic properties of human tendon structures in vivo. Eight subjects completed 8 weeks (4 days per week) of resistance training which consisted of unilateral plantar flexion at 70 % of one repetition maximum with 10 repetitions per set (5 sets per day). They performed resistance training (RT) on one side and resistance training and static stretching training (RST; 10 min per day, 7 days per week) on the other side. Before and after training, the elongation of the tendon structures in the medial gastrocnemius muscle was directly measured using ultrasonography, while the subjects performed ramp isometric plantar flexion up to the voluntary maximum, followed by a ramp relaxation. The relationship between estimated muscle force (F(m)) and tendon elongation (L) was fitted to a linear regression, the slope of which was defined as stiffness. The hysteresis was calculated as the ratio of the area within the F(m)-L loop to the area beneath the load portion of the curve. The stiffness increased significantly by 18.8 +/- 10.4 % for RT and 15.3 +/- 9.3 % for RST. There was no significant difference in the relative increase of stiffness between RT and RST. The hysteresis, on the other hand, decreased 17 +/- 20 % for RST, but was unchanged for RT. These results suggested that the resistance training increased the stiffness of tendon structures as well as muscle strength and size, and the stretching training affected the viscosity of tendon structures but not the elasticity.

Is passive stiffness in human muscles related to the elasticity of tendon structures?

The purpose of this study was to examine in vivo whether passive stiffness in human muscles was related to the elasticity of tendon structures and to performance during stretch-shortening cycle exercise. Passive torque of plantar flexor muscles was measured during passive stretch from 90 degrees (anatomical position) to 65 degrees of dorsiflexion at a constant velocity of 5 degrees.s-1. The slope of the linear portion of the passive torque-angle curve during stretching was defined as the passive stiffness of the muscle. The elongation of the tendon and aponeurosis of the medial gastrocnemius muscle (MG) was directly measured using ultrasonography during ramp isometric plantar flexion up to the voluntary maximum. The relationship between the estimated muscle force of MG and tendon elongation was fitted to a linear regression, the slope of which was defined as the stiffness of the tendon. In addition, the dynamic torques during maximal voluntary concentric plantar flexion with and without prior eccentric contraction were determined at a constant velocity of 120 degrees.s-1. There were no significant correlations between passive stiffness and either the tendon stiffness (r = 0.19, P > 0.05) or the relative increase in torque with prior eccentric contraction (r = -0.19, P > 0.05). However, tendon stiffness was negatively correlated to the relative increase in torque output (r = -0.42, P < 0.05). The present results suggested that passive stiffness was independent of the elasticity of tendon structures, and had no favourable effect on the muscle performance during stretch-shortening cycle exercise.

Effects of isometric training on the elasticity of human tendon structures in vivo.

The present study aimed to investigate the effect of isometric training on the elasticity of human tendon structures. Eight subjects completed 12 wk (4 days/wk) of isometric training that consisted of unilateral knee extension at 70% of maximal voluntary contraction (MVC) for 20 s per set (4 sets/day). Before and after training, the elongation of the tendon structures in the vastus lateralis muscle was directly measured using ultrasonography while the subjects performed ramp isometric knee extension up to MVC. The relationship between the estimated muscle force and tendon elongation (L) was fitted to a linear regression, the slope of which was defined as stiffness of the tendon structures. The training increased significantly the volume (7.6+/-4.3%) and MVC torque (33.9+/-14.4%) of quadriceps femoris muscle. The L values at force production levels beyond 550 N were significantly shorter after training. The stiffness increased significantly from 67.5+/-21.3 to 106.2+/-33.4 N/mm. Furthermore, the training significantly increased the rate of torque development (35.8 +/- 20.4%) and decreased electromechanical delay (-18.4+/-3.8%). Thus the present results indicate that isometric training increases the stiffness and Young's modulus of human tendon structures as well as muscle strength and size. This change in the tendon structures would be assumed to be an advantage for increasing the rate of torque development and shortening the electromechanical delay.

A closed-loop cadaveric foot and ankle loading model

Investigations of human foot and ankle biomechanics rely chiefly on cadaver experiments. The application of proper force magnitudes to the cadaver foot and ankle is essential to obtain valid biomechanical data. Data for external ground reaction forces are readily available from human motion analysis. However, determining appropriate forces for extrinsic foot and ankle muscles is more problematic. A common approach is the estimation of forces from muscle physiological cross-sectional areas and electromyographic data. We have developed a novel approach for loading the Achilles and posterior tibialis tendons that does not prescribe predetermined muscle forces. For our loading model, these muscle forces are determined experimentally using independent plantarflexion and inversion angle feedback control. The independent (input) parameters — calcaneus plantarflexion, calcaneus inversion, ground reaction forces, and peroneus forces — are specified. The dependent (output) parameters — Achilles force, posterior tibialis force, joint motion, and spring ligament strain — are functions of the independent parameters and the kinematics of the foot and ankle. We have investigated the performance of our model for a single, clinically relevant event during the gait cycle. The instantaneous external forces and foot orientation determined from human subjects in a motion analysis laboratory were simulated in vitro using closed-loop feedback control. Compared to muscle force estimates based on physiological cross-sectional area data and EMG activity at 40% of the gait cycle, the posterior tibialis force and Achilles force required when using position feedback control were greater.

Static and dynamic optimization solutions for gait are practically equivalent

The proposition that dynamic optimization provides better estimates of muscle forces during gait than static optimization is examined by comparing a dynamic solution with two static solutions. A 23-degree-of-freedom musculoskeletal model actuated by 54 Hill-type musculotendon units was used to simulate one cycle of normal gait. The dynamic problem was to find the muscle excitations which minimized metabolic energy per unit distance traveled, and which produced a repeatable gait cycle. In the dynamic problem, activation dynamics was described by a first-order differential equation. The joint moments predicted by the dynamic solution were used as input to the static problems. In each static problem, the problem was to find the muscle activations which minimized the sum of muscle activations squared, and which generated the joint moments input from the dynamic solution. In the first static problem, muscles were treated as ideal force generators; in the second, they were constrained by their force–length–velocity properties; and in both, activation dynamics was neglected. In terms of predicted muscle forces and joint contact forces, the dynamic and static solutions were remarkably similar. Also, activation dynamics and the force–length–velocity properties of muscle had little influence on the static solutions. Thus, for normal gait, if one can accurately solve the inverse dynamics problem and if one seeks only to estimate muscle forces, the use of dynamic optimization rather than static optimization is currently not justified. Scenarios in which the use of dynamic optimization is justified are suggested.

Influence of static stretching on viscoelastic properties of human tendon structures in vivo.

The purpose of this study was to investigate the influences of static stretching on the viscoelastic properties of human tendon structures in vivo. Seven male subjects performed static stretching in which the ankle was passively flexed to 35 degrees of dorsiflexion and remained stationary for 10 min. Before and after the stretching, the elongation of the tendon and aponeurosis of medial gastrocnemius muscle (MG) was directly measured by ultrasonography while the subjects performed ramp isometric plantar flexion up to the maximum voluntary contraction (MVC), followed by a ramp relaxation. The relationship between the estimated muscle force (Fm) of MG and tendon elongation (L) during the ascending phase was fitted to a linear regression, the slope of which was defined as stiffness of the tendon structures. The percentage of the area within the Fm-L loop to the area beneath the curve during the ascending phase was calculated as an index representing hysteresis. Stretching produced no significant change in MVC but significantly decreased stiffness and hysteresis from 22.9 +/- 5.8 to 20.6 +/- 4.6 N/mm and from 20.6 +/- 8.8 to 13.5 +/- 7.6%, respectively. The present results suggest that stretching decreased the viscosity of tendon structures but increased the elasticity.

Estimation of muscle forces in the lumbar spine during upper-body inclination

Objective. To estimate the muscle forces during upper-body inclination and to determine their influence on stress distribution in the annulus fibrosus of the lumbar spine discs. Design. The muscle forces and stresses were calculated using a non-linear finite element model of the lumbar spine. Background. Little is known about the influence of muscle forces on the deformation of, and stresses in, the lumbar spine. In most studies, muscle forces are neglected. Methods. Three-dimensional non-linear finite element models of the ligamentous lumbar spine, with and without internal spinal fixators, were created. They were validated by use of experimental data from in vitro measurements on cadaver specimens. In a second step, the influence of muscle forces on stresses in the annulus fibrosus of the lumbar spine discs was investigated in a parameter study. This was done for different inclination angles of the upper-body. Results. Good agreement between analytical and experimental results proved achievable when loading with pure moments in the three main planes of the lumbar spine. For inclination of the upper-body, the flexion angle clearly has a strong influence on the stresses in the lumbar spine while the influence of local muscles was small. The stress distribution in the discs differed considerably when the muscle forces are neglected and only a pure moment is applied. Conclusions. This study confirmed earlier ones that have shown that muscle forces should not be neglected when studying the stresses in the lumbar spine. The local dorsal muscles, however, have only a small influence on the stresses in the discs. Relevance For investigations of the biomechanical effects of spinal implants and surgical procedures, experimental or analytical methods are used. Due to the complexity involved, as well as to a lack of information, muscle forces are often neglected. Our study showed that muscles do in fact have a major influence on the mechanical behaviour of the spine and should always be taken into account.

Questions to ask When Interpreting Surface Electromyography (SEMG) Research

Surface electromyography (SEMG) is widely used to evaluate muscle activity. In SEMG, researchers attach electrodes to the surface of the skin overlying a muscle and measure the amount of electricity it produces as muscle fibers contract. SEMG can determine which muscles are active, their degree of activity, and how active the muscle is compared to the subject's capacity. It can also be used to estimate muscle force. Properly employed, SEMG assists in evaluating the relative risk of a work task. As articles reporting SEMG results are often used by ergonomics practitioners as guidance in job design, the ability to interpret SEMG research is critical. Problems occur when researchers assume their readers have a greater familiarity with SEMG than actually exists, or when they make any of a number of SEMG-related research or interpretation errors. This paper suggests some questions that should be asked when evaluating a study that reports SEMG data.

Prediction of muscle force involved in shoulder internal rotation

The estimation of shoulder muscle force is important to understand the mechanism of rotator cuff injury. In general, mean parameter values are used as input for computational models. However, anatomic and biomechanical parameters vary widely among people. The purpose of this study was to evaluate shoulder muscle forces predicted by an electromyography-driven muscle model given neuromuscular parameters generated by Monte Carlo simulation. Normal distributions were used to model muscle moment arms; electromyographic and muscle physiological cross-sectional area data were modeled with log-normal distributions. Eight muscles were included in the model. Muscle force and joint moment were predicted on the basis of the simulated parameters. The results showed that the subscapularis and pectoralis major were substantial actuators for shoulder internal rotation. During maximum voluntary contraction the median of the muscle forces of the subscapularis and the pectoralis major were 1030 N and 462 N, respectively. This study demonstrated that the Monte Carlo method could be used for muscle force prediction by integrating population variability of physiological parameter into a biomechanical muscle model.

Prediction of muscle force involved in shoulder internal rotation

The estimation of shoulder muscle force is important to understand the mechanism of rotator cuff injury. In general, mean parameter values are used as input for computational models. However, anatomic and biomechanical parameters vary widely among people. The purpose of this study was to evaluate shoulder muscle forces predicted by an electromyography-driven muscle model given neuromuscular parameters generated by Monte Carlo simulation. Normal distributions were used to model muscle moment arms; electromyographic and muscle physiological cross-sectional area data were modeled with log-normal distributions. Eight muscles were included in the model. Muscle force and joint moment were predicted on the basis of the simulated parameters. The results showed that the subscapularis and pectoralis major were substantial actuators for shoulder internal rotation. During maximum voluntary contraction the median of the muscle forces of the subscapularis and the pectoralis major were 1030 N and 462 N, respectively. This study demonstrated that the Monte Carlo method could be used for muscle force prediction by integrating population variability of physiological parameter into a biomechanical muscle model.

Muscle forces and spinal loads at C4/5 level during isometric voluntary efforts.

The goal of this study was to determine neck muscle forces and spinal loads that result from isometric muscle contractions. Electromyographic (EMG) activity of the neck musculature and a three-dimensional biomechanical model of the neck were used. The model was EMG-based and estimated muscle forces and spinal loads at the C4/5 level. EMG signals were collected from eight sites at the C4/5 level of the neck using Ag-AgCl surface electrodes from 10 adult male subjects. The subjects performed isometric contractions gradually developing to maximum efforts in flexion, extension, left lateral bending, and right lateral bending. During maximum voluntary contraction (MVC) trials most muscles generated high levels of EMG signal during cervical rotation. The posterior surface of the neck (trapezius) was the only electrode site at which maximum activity EMG consistently occurred by the same method (rotation) in all subjects. Variations in the EMG patterns were observed in different experiments that produced overall neck moments of equal magnitudes. With these data the model computed variations in load distribution among the agonist muscles. Consistent also with EMG distributions, the model also computed co-contractions of antagonist muscles. The average (+/- SD) magnitudes of peak moments were 28.3 (+/- 3.3) Nm in extension, 17.7 (+/- 3.1) Nm in flexion, 16.9 (+/- 2.8) Nm in left lateral bending, and 17.0 (+/- 2.9) Nm in right lateral bending. The model predicted C4/5 joint compressive forces during peak moments were 1372 (+/- 140) N in extension, 1654 (+/- 308) N in flexion, 956 (+/- 169) N in left lateral bending, and 1065 (+/- 207) N in right lateral bending. Results suggest that higher C4/5 joint loads than previously reported are possible during maximum isometric muscle contractions.

Macaque Pterygoid Muscles: Internal Architecture, Fiber length, and Cross-Sectional Area

Models of mastication require knowledge of fiber lengths and physiological cross-sectional area (PCS), a proxy for muscle force. I dissected 36 medial pterygoid and 36 lateral pterygoid muscles from 30 adult females of 3 macaque species (Macaca fascicularis, M. mulatta, M. nemestrina) using gross and chemical techniques and calculated PCS. These macaques have mechanically similar dietary niches and exhibit no significant difference in muscle architecture or fiber length. Fiber length does not scale with body size (mass) for either total pterygoid muscle or for medial pterygoid muscle mass. However, fiber length scales weakly with lateral pterygoid muscle mass. In each case, differences in PCS among species result from differences in muscle mass not fiber length. Medial pterygoid PCS scales isometrically with body size; larger animals have greater force production capabilities. Medial and lateral pterygoid PCS scale positively allometrically with facial size; individuals with more prognathic faces and taller mandibular corpora have greater PCS, and hence force, values. This positive allometry counters the less efficient positioning of masticatory muscles in longer-faced macaques. PCS is only weakly correlated with bone proxies previously used to estimate muscle force. Thus, predictions of muscle force from bone parameters will entail large margins of error and should be used with caution.

Difficulties in estimating muscle forces from muscle cross-sectional area. An example using the psoas major muscle.

Most biomechanical models use muscle cross-sectional area (CSA) as an indicator of maximum isometric muscle force. In general, there are multiple estimates of CSA for the same muscle. For example, numerous studies have estimated the CSA of the psoas major muscle using different subject populations and positions. However, few studies have combined the available information to obtain an overall estimate of CSA or investigated the effect different subject characteristics may have on CSA. In the present update, nine studies that reported psoas major CSA or physiologic CSA were compared with respect to subject characteristics, methodology, and results. Corrections to cadaveric data were made to adjust physiologic CSA to CSA. Comparison of reported values for living subjects indicated that females have smaller mean CSA than males for the psoas major muscle and that body size does not significantly influence muscle CSA in males. Areas derived from cadaveric data were smaller than similar studies on living subjects, possibly because of subject age, removal of tendinous and fatty components of fascicles, and lack of detailed data for fascicle angles in the supine position. Results indicate that researchers who use muscle CSA in biomechanical models should carefully assess the appropriateness of the data used, particularly in relation to potential sex differences and the influence of postural changes on CSA.

Estimation of trunk muscle forces using the finite element method and in vivo loads measured by telemeterized internal spinal fixation devices

Direct quantitative measurement of muscle forces is not possible. Forces in the trunk muscles were estimated for standing and flexion of the upper body using three-dimensional, nonlinear finite element models of the lumbar spine with and without an internal spinal fixation device. Muscle forces assumed were two pairs dorsally and one ventrally, each representing several muscles. Muscle forces in the model with internal fixators were varied in discrete steps until the implant loads calculated closely corresponded to those measured in a patient with an instrumented implant. The calculated angles between adjacent lumbar vertebrae were compared with corresponding values measured on X-ray films of a patient as well as with literature values and served as a second criterion for predicting muscle forces. For the model without an implant, the muscle forces of the first model were slightly varied until the lumbar spine shape and the intradiscal pressure were physiological. The abdomen was shown to have a considerable supporting function for flexion.

Total trunk muscle force and spinal compression are lower in asymmetric moments as compared to pure extension moments

The aim of the present study was to test the assumption that asymmetric trunk loading requires a higher total muscle force and consequently entails a higher compression forces on the spine as compared to symmetric loading. When the trunk musculature is modelled in sufficient detail, optimisation shows that there is no mechanical necessity for an increase in total muscle force (or compression force) with task asymmetry. A physiologically based optimisation does also not predict an increase in total muscle force or spinal loading with asymmetry. EMG data on 14 trunk muscles collected in eight subjects showed antagonistic coactivity to be present in both conditions. However, estimates of total muscle force based on the EMG were lower when producing an asymmetric moment. In conclusion, producing an asymmetric moment appears to cause slightly lower forces on the lumbosacral joint as compared to a symmetric moment. Only lateral shear forces increase with asymmetry but these remain well below failure levels.

ヒト生体における腓腹筋内側頭の腱組織の粘弾性

The purpose of this study was to investigate the viscoelastic properties of tendon structures in humans. Elongation of the tendon and aponeurosis of medial gastrocnemius muscle (MG) was directly measured by ultrasonography, while subjects (N=12) performed ramp isometric plantar flexion up to the voluntary maximum, followed by a ramp relaxation. The relationship between estimated muscle force (Ff) and tendon elongation (dL) was fitted to a linear regression, the slope of which was defined as stiffness of the tendon structures. The hysteresis was calculated as the ratio of the area within the Ff-dL loop (elastic energy dissipated) to the area beneath the load portion of the curve (elastic energy input) . The resulting Ff-dL relationship was non-linear in form, as previously reported on animal and human tendons in vitro. The mean stiffness was 24.0±5.6 N/mm. However, there was a considerable inter-subject variability (15.8 to 36.8 N/mm) . The Young's modulus, i. e., the slope of the stress-strain curve, was 280 MPa, which tended to be lower than the previously reported values for human tendons. It was also found that the strain of the tendon structures was homogeneously distributed along its length. The mean hysteresis (energy dissipation) was 23.4±12.4%. However, again there was a considerable inter-subject variability (8.7 to 39.3%) . The present results indicated that the tendon structures of human MG was considerably compliant and its hysteresis was in accordance with previously reported values.

Macaque Masseter Muscle: Internal Architecture, Fiber Length and Cross-Sectional Area

Models of mastication require knowledge of fiber lengths and physiological cross-sectional area (PCS): a proxy for muscle force. Yet only a small number of macaques of various species, ages, and sexes inform the previous standards for masseter muscle architecture. I dissected 36 masseters from 30 adult females of 3 macaque species—Macaca fascicularis, M. mulatta, M. nemestrina—using gross and chemical techniques and calculated PCS. These macaques have mechanically similar dietary niches and exhibit no significant difference in masseter architecture or fiber length. Intramuscular tendons effectively compartmentalize macaque masseters from medial to lateral. Fiber lengths vary by muscle subsection but are relatively conservative among species. Fiber length does not scale with body size (mass) or masseter muscle mass. However, PCS scales isometrically with body size; larger animals have greater force production capabilities. PCS scales positively allometrically with facial size; animals with more prognathic faces and taller mandibular corpora have greater PCS, and hence force, values. This positive allometry counters the less efficient positioning of masseter muscles in longer-faced animals. In each case, differences in PCS among species result from differences in muscle mass not fiber length. Masseter PCS is only weakly correlated with bone proxies previously used to estimate muscle force. Thus predictions of muscle force from bone parameters will entail large margins of errors and should be used with caution.

Estimation of trunk muscle forces and spinal loads during fatiguing repetitive trunk exertions.

The effects of human trunk extensor muscle fatigue on the estimated trunk muscle forces and spinal loading were investigated during the performance of repetitive dynamic trunk extension. To evaluate if alterations in the trunk muscle recruitment patterns resulted in a greater estimated active loading of the spine and, in turn, an increased risk of injury. Epidemiologic studies highlight the increased risk of low back injury during repetitive lifting, implicating fatigue of muscles and/or passive tissues as causes of such injury. Increased trunk muscle activity or altered recruitment patterns resulting from fatigue in the primary trunk extensor muscles may indicate an increase in the active loading of the spine, which could contribute to an increased risk of injury. Sixteen healthy study participants performed repetitive isokinetic trunk extension endurance tests at two load levels and two repetition rates, while their net muscular torque output and trunk muscular activity were measured. During each exertion, trunk torque, position, and velocity were controlled, so that any change in muscle activity could be attributed to fatigue. An electromyography-assisted model, adapted to accommodate the decline in maximum muscular tension generation resulting from fatigue, was used to estimate the 10 trunk muscle forces and spinal loading. Linear regression was used to quantify the rate of change in muscle force and spinal loading resulting from fatigue, while analysis of variance was used to determine if the rate of change was dependent on the task conditions (load and repetition rate). Significant elevations were estimated for the latissimus dorsi and external oblique muscle forces in more than 70% of the endurance tests, whereas significant reductions in the erector spinae muscle force were predicted in 75% of the trials. The magnitude of the range of change of the erector spinae and latissimus dorsi muscle forces was dependent on the load level and repetition rate. The reduction in erector spinae forces offset the augmented force in the other muscles, because the net changes in compression and lateral shear forces on the spine were not significant, and the anteroposterior shear was reduced. The results of the study do not suggest that an increase in the muscular loading of the spine occurs as a result of changing trunk muscular recruitment patterns. Therefore, future studies should focus on injury mechanisms that may occur as a result of a change in the viscoelastic passive tissue responses, muscular insufficiency, or a decline in neuromuscular control and coordination.