Accommodative foot orthoses are used mainly to provide cushioning and shock absorption and thus improved stability and forward propulsion. Insoles are fitted to subjects with pathological gait as well as to normal individuals for use during daily and/or sporting activities. The range of available ready-made insoles has broadened, but at the same time the selection of the most appropriate device for each individual has become more difficult.
Several studies have assessed the properties of the materials used to fabricate mass-produced and/or custom-made insoles, 1–4 the clinical effectiveness of treatment of abnormal states using various foot orthoses, 5,6 and the correlation of orthoses use with specific athletic injuries. 7,8
However, research on the biomechanical effectiveness of foot orthoses in normal individuals is limited and is focused on the effect of insoles on the kinematics of the lower limbs during running 9,10 and on plantar pressure redistribution. 11–13 Research on the effect of foot orthoses on pelvis and trunk kinematics is scarce. 14
In the current study, we investigated how normal subjects respond, in terms of the motion of their lower limbs, pelvis, and trunk, during level walking when fitted with mass-produced separated medial and lateral arms insoles.
MATERIALS AND METHODS
The foot orthosis used in this study was a three-quarters length, U-shaped insole with separated medial and lateral arms, made of expanded closed-cell foam (Figure 1). This orthosis possesses a deep offset heel seat to cup the calcaneus. High extended and separated arms were formed to provide independent column control. The maximum thickness of the insoles used was 5.2 mm.
All subjects wore the orthoses in conjunction with personal leisure shoes. The shoes had adequate space to accommodate and complement the added material. All shoes had a full-length flat sole made of rubber, and they were closed heel with a heel height difference (heel height minus sole height) between 13 and 15 mm.
Subjects were tested under three conditions: barefoot, shod, and shod with insoles. The dependent variables were kinematic gait parameters.
Gait analysis was conducted during two separate sessions. During the first session, subjects were tested while barefoot and while wearing shoes. Insoles were worn during the second session, held 3 weeks after the first to achieve a minimum degree of acclimation to the insoles. The fitting of the insoles was carried out in a consistent manner by the same person.
Eight asymptomatic, healthy young male subjects participated in this study. Inclusion criteria were: (1) no current musculoskeletal system complications; (2) right-handed and right-footed; (3) absence of any cognitive problems; (4) absence of any other conditions that could affect and/or limit walking ability, and (5) wearing the insoles for 3 weeks preceding the second gait analysis session.
The foot structure of the participants was characterized as “neutrally” aligned, with no deformities that could justify a pes planus or pes cavus type. None had experienced any neurologic or muscular disorder or injury to the tibialis posterior tendon, or any condition in which one or more of the arches of the foot had fallen. Furthermore, no foot deformities that might lead to abnormal positions of the calcaneus were revealed under static and dynamic observation. Finally, there were no other signs, such as muscle atrophy, bony prominences, or toe and metatarsal head deformities, that could be related to the presence of foot abnormalities.
The measuring procedure was carried out according to the ethical guidelines for teaching and research edited by the Advisory Committee on Ethics of the University of Surrey, and informed consent was obtained from all participants.
The mean age of the subjects was 24.4 ± 4.3 years, mean weight 64.5 ± 8.3 kg, and mean height 1.77 ± 0.08 m.
Three-dimensional (3-D) gait analysis, using seven CCD cameras (ProReflex MCU, Qualisys Medical AB, Gothenburg, Sweden) and passive retroreflective skin markers, was performed on all participants. Before any image capture, the system was calibrated using a reference structure for defining the calibration coordinate system and a wand to provide the camera system with measurement points. The distance between two 19.3 mm markers affixed to a rigid rod (wand) such that their centers were 750 mm apart, was recorded as the wand was moving in all three directions and in the entire measurement area within a space of 4000 × 2000 × 2000 mm. During all calibration measurements, the standard deviation was less than 0.5% and the range of difference less than 1%.
The participants were asked to wear close-fitting, dark-colored swimsuits made of elastic, nonreflective material. Seven spherical retroreflective markers with a diameter of 19.3 mm were placed on both sides of their lower body at specific anatomical points: between the second and third metatarsal bones (15 mm proximally of the metatarsal heads); on the heel (posterior of the calcaneus, at the same horizontal plane as the previous marker); on the lateral malleolus; inferior to the knee (on the middle of the tuberosity of tibia); on the knee joint line; superior to the knee (10 mm superior to the patella), and on the anterior superior iliac spine. In addition, another set of four markers was placed on the trunk and pelvis: on the sacrum, on the 12th thoracic vertebrae, and on both acromion (Figure 2).
The skin at the sites of marker placement was cleaned with an alcohol wipe to remove any oils or lotions. The markers were then placed with hypoallergenic tape. The placement of the skin-mounted markers was carried out in a consistent manner by the same person.
During each session and before actual image capturing, an orientation period allowed the participants to practice walking under testing conditions. Subjects were asked to walk on the 10-m walkway four times to determine a comfortable speed. The participant’s performance during six consecutive gait cycles was then captured at the same rate.
To compensate for nonpathologic variations and anatomical differences among the participants, a static angular offset measurement was carried out before each gait trial. Subjects were asked to stand in the middle of the calibrated measurement area in a relaxed position. Measurements were recorded for 15 seconds, at a frequency of 240 Hz. All subjects started from a position 4 m away from the measurement area to reach a natural continuous walking pattern once they entered the calibrated area.
The captured images were simultaneously digitized, and all 3-D segments were processed and edited using the QTrac software (Qualisys Medical AB). The procedure included tracking motion data (creating 3-D marker trajectories); sorting the 3-D data according to the markers used in the measurement, and selecting appropriate data to export for analysis.
To remove any additive noise (ie, any components of the final signal that were not attributable to the process itself), the collected raw data for the Cartesian coordinates of the 18 markers during standing and walking were smoothed by means of a second order low-pass Butterworth filter, applied both backward and forward, resulting in a fourth order filter, with a cutoff frequency of 10 Hz.
Using the filtered data, we calculated the angles (in degrees) of the hip, knee, ankle-foot, pelvis, and trunk using QGait software (Qualisys Medical AB). The standing trial angles, used to derive the neutral position reference, were then subtracted from the gait trial angles.
Several spatial variables were derived from the angular position data of the right leg, the pelvis, and the trunk. The ankle-foot data at the frontal plane and the hip joint data at the transverse plane were not used because of the high level of noise associated with them, which was present despite the filtering. The noise may be attributed to problems related to the small distances between the markers in the area of the ankle-foot complex, making it difficult for the cameras to differentiate between markers.
The angular kinematic data from each of the six recorded gait cycles for each subject were averaged. Data for all eight subjects were then averaged to produce a single data set representing the entire sample.
The angles at the sagittal, frontal, and transverse planes were determined as follows: (1) hip: flexion/extension, adduction/abduction, and internal rotation/external rotation (not used); (2) knee: flexion/extension, adduction (varus)/abduction (valgus), and internal rotation/external rotation; (3) ankle-foot: dorsiflexion/plantarflexion, inversion/eversion (not used), and internal rotation/external rotation; (4) pelvis: anterior tilt/posterior tilt, up obliquity/down obliquity, internal rotation/external rotation; and (5) trunk: flexion/extension, lateral bend/medial bend, and retraction/protraction.
The gait cycle duration (in seconds), the stride length (in meters), the walking speed (in meters per second), the cadence (in steps per minute), and the right limb stance phase duration (in seconds) were also calculated.
One-way analysis of variance (ANOVA) tests were performed to characterize the differences of the measured parameters in the three experimental conditions, and the Tukey-Kramer multiple comparisons tests were used when significant differences were found (p < .05).
Typical diagrams of the angular position of the hip, knee, ankle-foot, pelvis, and trunk during a complete gait cycle for each test condition in the sagittal plane are presented in Figure 3. Figure 4 presents the same data for the hip, knee, pelvis and trunk;Figure 5, for the knee, ankle-foot, pelvis and trunk.
The angular variables for which statistically significant differences were observed are presented in Table 1.
Observed differences in hip and knee flexion and extension, and angle at heel strike were not statistically significant.
Maximum dorsiflexion in the ankle joint significantly increased by 28.5% when the insole condition was compared with the barefoot condition. Maximum plantarflexion decreased but not significantly. Although subjects 6 and 7 showed high values of dorsiflexion influencing the mean value of the sample, all other subjects showed consistently high maximum dorsiflexion when the insoles were used. The angle at heel strike significantly decreased when both the shod and the insole conditions were compared with the barefoot condition, by 560% and 54.3%, respectively. In addition, all subjects when fitted with the orthosis showed consistently higher heel strike angle than that observed with the shoes alone. The first peak in plantarflexion was increased when the insoles were used. The highest difference (47.9%) was observed when the shod condition was compared with the insole condition.
No significant differences were found in the pelvis or the trunk for any of the tested variables.
When the insoles were used, hip angle at heel strike decreased by 84.6% and 56.0%, when the insole condition was compared with the barefoot and shod conditions, respectively. The observed difference was statistically significant when the insole and shod conditions were compared with the barefoot condition.
No significant differences were found in the knee joint or pelvis for any of the analyzed variables. In the trunk, when the insoles were used, the maximum lateral bend was significantly reduced by 83.8% and 86.1% when compared with the barefoot and shod conditions, respectively. Furthermore, for all subjects, the lowest value of the maximum lateral bend angle was observed when the insoles were used, and the highest one in the shod condition.
Maximum internal rotation angle in the ankle joint significantly decreased by 52.7% and 48.7% when the insole condition was compared with the barefoot and shod conditions, respectively. In all cases, the highest value of the maximum internal rotation angle was observed under the shod condition. The maximum external rotation angle also decreased but not significantly. The angle at heel strike significantly decreased by 52.6% when the insole condition was compared with the shod condition. Although subject 7 showed a high absolute heel strike angle influencing the mean value of the sample, all other subjects showed consistently high values when the shoes were used. In addition, all subjects but one (subject 3) exhibited their lowest absolute value when fitted with the insoles. The first peak in external rotation was decreased but not significantly.
No significant differences were found in the pelvis or trunk for any of the tested variables.
The gait variables for which statistically significant differences were observed–cycle time and cadence–are presented in Table 2.
When the insoles were used, the duration of a complete gait cycle decreased by 1.3% when compared with the shod condition but increased by 5.9% when compared with the barefoot condition. All subjects exhibited their lowest value when tested barefoot and their highest value when shod (with the exception of subject 2). The observed differences were statistically significant because of the small p value observed when the barefoot and shod conditions were compared.
When the insoles were used, the average cadence of the subjects decreased by 5.5% when compared with the barefoot condition and increased by 1.3% when compared with the shod condition. The observed differences were statistically significant because of the small p value observed when the barefoot and shod conditions were compared. In addition, all subjects except subject 2 showed an increased cadence when changing from the shod to the insole condition.
Observed differences in stride length, walking speed, and stance time were not statistically significant.
Averaging procedures, to minimize the effects of the intrinsic variability of human gait, were applied to determine the mean values and standard deviations of the measured parameters at each test condition. Thus, reported differences refer to the absolute mean values of a single data set of the entire sample. However, they seem to represent quite well the trend observed in each participant individually.
Our results showed that statistically significant changes caused by the use of the separated-arms insoles were observed mainly for the ankle-foot complex at the sagittal and transverse planes. All subjects showed consistently higher values of dorsiflexion angle at heel strike and maximum dorsiflexion angle when the insole condition was compared with the shod one. This was not the case for the external rotation angle at heel strike and the maximum internal rotation angle, which were both decreased. Most of the angular position changes at the hip, knee, pelvis, and trunk were not statistically significant, the exceptions being the hip adduction at heel strike and the trunk maximum lateral bend. All subjects showed the lowest value of their maximum lateral bend angle when fitted with the insoles. In addition, all of them had a consistently lower hip adduction angle at heel strike, when the insole condition was compared with the barefoot condition. Alterations of the hip joint and trunk motion were found at the sagittal plane; these alterations, although not statistically significant, exhibited p values close to the set limit of significance. In addition, changes of the knee joint motion parameters were observed, mainly at the sagittal and transverse planes.
The maximum anterior and posterior tilt of the pelvis and the maximum internal and external rotation angles all decreased when the insoles were used. In general, all the pelvic variables decreased, except for the maximum down obliquity and the heel strike angle at the frontal plane. Although the observed differences were not statistically significant, they are of some clinical importance considering the high cost of human locomotion.
When trunk motion was considered, the maximum flexion and extension angles both increased, whereas the maximum retraction and protraction angles decreased, with the differences being more prominent in the cases of the maximum flexion and the maximum retraction angles.
The separated-arms orthoses might affect the motion of the ankle-foot complex at the frontal plane and the hip joint at the transverse plane; however, this was not investigated because the corresponding data could not be used.
In general, although some noticeable changes were observed for the ankle, knee, hip and trunk, the pelvis was not significantly affected by the study orthoses. The reason for this remains unclear and indicates that foot orthoses might also affect the soft tissues of the lower body and the neurologic control of gait, as suggested by others. 14
From the analysis of the most relevant gait parameters (ie, the linear kinematics parameters), statistically significant differences were found only for cycle time and cadence. In addition, post-tests revealed that these were attributable to the use of shoes.
All subjects exhibited consistently lower gait cycle time, higher stride length and walking speed, and slightly higher cadence when changing from the shod to the insole condition. Furthermore, all subjects but one showed a decrement in their stance phase duration. These findings might be of some clinical importance because the associated parameters are commonly used to evaluate the functional outcome of any orthotic intervention.
As reported, subjects were allowed 3 weeks to achieve a minimum degree of acclimation to the insoles. Although, as shown by our study, there are immediate changes to the walking pattern because of the separated-arms orthoses, the maximal effects may be apparent only after several weeks or months. In addition, only level walking was considered, another limitation of our study.
This study showed that, for the examined group of healthy individuals with the specific demographic characteristics, there were some systematic differences resulting from the short-term use of the separated-arms foot orthoses, mainly observed in the ankle-foot complex.
Alterations of the knee, hip, pelvis and trunk motion were found; these alterations, although not statistically significant, exhibited p values very close to the set limit of significance.
Most subjects exhibited consistently lower gait cycle and stance phase duration and higher stride length, walking speed, and cadence when changing from the shod to the insole condition.
The effects these insoles have on the soft tissues of the lower body and trunk, and the neurologic control of gait, which in conjunction with the observed kinematic changes, might result in pain and fatigue relief, remain unclear.