lower limb movement. We used functional magnetic resonance imaging (fMRI) to examine brain activation during pedaling in people with and without stroke ( Promjunyakul, Schmit, & Schindler-Ivens, 2015 ). People with stroke displayed reduced pedaling-related brain activation volume compared with age
Brice T. Cleland and Sheila Schindler-Ivens
James Wright, Thomas Walker, Scott Burnet and Simon A. Jobson
recently pedal-based systems have not provided the same measure of reliability when compared with more traditional crank- or hub-based systems, with Sparks et al 6 suggesting that the Look Kéo power pedals were not as reliable as the SRM Powermeter during an incremental testing protocol. Recently, the
Jeffrey B. Wheeler, Robert J. Gregor and Jeffrey P. Broker
In response to the popularity of clipless bicycle pedals with float designs, an instrumented force pedal system with multicompatibility for different shoe/pedal interfaces is presented. A dual piezoelectric element pedal has been modified for use with popular clipless pedal interfaces. The dual transducer arrangement permits measurement of three components of uniaxial load, location of the applied load, and calculation of the moment Mz about an axis through the position of the applied load and orthogonal to the pedal surface. Quantification of lower extremity kinetics using float feature pedals and the investigation of the pathomechanics of lower extremity cycling overuse injuries, especially knee injuries, is warranted. Qualitative descriptions of lower extremity pathomechanics related to overuse injuries have suggested that foot constraint may induce undesirable knee kinematics and kinetics. The instrumented force pedal system described here permits a comparison between pedal kinematics and kinetics of popular shoe/pedal interfaces with varying degrees of float allowance.
Graham E. Caldwell, Li Li, Steve D. McCole and James M. Hagberg
Alterations in kinetic patterns of pedal force and crank torque due to changes in surface grade (level vs. 8% uphill) and posture (seated vs. standing) were investigated during cycling on a computerized ergometer. Kinematic data from a planar cine analysis and force data from a pedal instrumented with piezoelectric crystals were recorded from multiple trials of 8 elite cyclists. These measures were used to calculate pedal force, pedal orientation, and crank torque profiles as a function of crank angle in three conditions: seated level, seated uphill, and standing uphill. The change in surface grade from level to 8% uphill resulted in a shift in pedal angle (toe up) and a moderately higher peak crank torque, due at least in part to a reduction in the cycling cadence. However, the overall patterns of pedal and crank kinetics were similar in the two seated conditions. In contrast, the alteration in posture from sitting to standing on the hill permitted the subjects to produce different patterns of pedal and crank kinetics, characterized by significantly higher peak pedal force and crank torque that occurred much later in the downstroke. These kinetic changes were associated with modified pedal orientation (toe down) throughout the crank cycle. Further, the kinetic changes were linked to altered nonmuscular (gravitational and inertial) contributions to the applied pedal force, caused by the removal of the saddle as a base of support.
Håvard Lorås, Gertjan Ettema and Stig Leirdal
Changes in pedaling rate during cycling have been found to alter the pedal forces. Especially, the force effectiveness is reduced when pedaling rate is elevated. However, previous findings related to the muscular force component indicate strong preferences for certain force directions. Furthermore, inertial forces (due to limb inertia) generated at the pedal increase with elevated pedaling rate. It is not known how pedaling rate alters the inertia component and subsequently force effectiveness. With this in mind, we studied the effect of pedal rate on the direction of the muscle component, quantified with force effectiveness. Cycle kinetics were recorded for ten male competitive cyclists at five cadences (60–100 rpm) during unloaded cycling (to measure inertia) and at a submaximal load (~260 W). The force effectiveness decreased as a response to increased pedaling rate, but subtracting inertia eliminated this effect. This indicates consistent direction of the muscle component of the foot force.
Harsh H. Buddhadev, Daniel L. Crisafulli, David N. Suprak and Jun G. San Juan
range of motion. 10 To our knowledge, however, no research has examined pedaling mechanics bilaterally in individuals with knee OA during cycling under submaximal conditions, despite the known therapeutic benefits of cycling. During cycling, power output at the crank is representative of net muscular
Birgit Larsen, Michael Voigt and Michael J. Grey
The influence of pedaling frequency and crank load on the sensitivity of the soleus short latency stretch reflex (SLR) was examined in nine healthy subjects during pedaling by the use of a custom-built robotic actuator. The SLR decreased successively in downstroke when pedaling frequency increased from 20 to 40 and 60 revolutions per minute at a constant crank load (p = .005). The SLR was unchanged at crank load increases of 2.6 or 5.1 Nm at a constant pedaling frequency (p > .05). Accordingly, it was shown that increased muscle activation level as a consequence of added crank load and increased movement speed does not increase the sensitivity of the soleus SLR.
Brian R. Umberger and Philip E. Martin
Lower extremity motions during cycling are often assumed to occur in the sagittal plane. While seemingly logical, this assumption has not been rigorously tested. Frontal plane rotation of the ankle joint (inversion/eversion) has been studied extensively during gait but infrequently during cycling despite the suggestion that excessive eversion or pronation may be related to overuse knee injuries. Two-dimensional sagittal plane hip, knee, and ankle joint kinematics were generally found to be similar to simultaneously measured 3-D values. Despite the similarity in motion patterns, maximum hip angle was 34° more flexed in 2-D than 3-D. Maximum and minimum frontal plane ankle joint angles were similar in 2-D and 3-D. However, during the middle of the pedal cycle, 2-D frontal plane ankle joint motion deviated from 3-D, such that maximum ankle eversion was reached 36% of the pedal cycle later in 2-D versus 3-D. The discrepancy at the hip was due primarily to differences in hip angle definition for 2-D and 3-D approaches, and an alternate convention for hip angle in 2-D is suggested. Discrepancies in frontal plane ankle joint motion are due to weaknesses in the planar approach and would be difficult to overcome without resorting to 3-D measurement.
Cheryl D. Pierson-Carey, David A. Brown and Christine A. Dairaghi
The purpose of this study was to determine the effects of limiting ankle motion on pedal forces. Sixteen adults pedaled an instrumented ergometer against constant cadence and frictional load while wearing hinged braces. Ankle motion was limited under four randomly assigned conditions: both braces unlocked (UL), only the preferred leg (PL) brace locked, only the nonpreferred leg (NPL) brace locked, and braces on both legs (BL) locked. Measurements of pedal force, crank, and pedal angles were sampled at 200/s for 20 s. With both braces locked, resultant force mean magnitude decreased during the downstroke, due to reduced radial crank force. Asymmetry between PL and NPL decreased during the power phase when only PL was braced but increased when only NPL was braced. It was concluded that constrained ankle motion, as may occur with ankle injury or hemiplegia, reduces the ability to transmit power during the downstroke while enhancing ability during the upstroke.
Kreg G. Gruben, Lynn M. Rogers, Matthew W. Schmidt and Liming Tan
The force that healthy humans generated against a fixed pedal was measured and compared with that predicted by four models. The participants (n = 11) were seated on a stationary bicycle and performed brief pushing efforts against an instrumented pedal with the crank fixed. Pushes were performed to 10 force magnitude targets and at 12 crank angles. The increasing magnitude portion of the sagittal-plane force path for each push effort was fitted with a line to determine the direction of the muscle component of the foot force. Those directions varied systematically with the position of the pedal (crank angle) such that the force path lines intersected a common region superior and slightly anterior to the hip. The ability of four models to predict force path direction was tested. All four models captured the general variation of direction with pedal position. Two of the models provided the best performance. One was a musculoskeletal model consisting of nine muscles with parameters adjusted to provide the best possible ft. The other model was derived from (a) observations that the lines-of-action of the muscle component of foot force tended to intersect in a common region near the hip, and (b) the corresponding need for foot force to intersect the center-of-mass during walking. Thus, this model predicted force direction at each pedal position as that of a line intersecting the pedal pivot and a common point located near the hip (divergent point). The results suggest that the control strategy employed in this seated pushing task reflects the extensive experience of the leg in directing force appropriately to maintain upright posture and that relative muscle strengths have adapted to that pattern of typical activation.