Abstract: The design of high performance cycling shoes is based upon an understanding of the nature of the demands to be placed upon that shoe during its lifetime. Demands can be viewed as a direct consequence of the forces applied by the cyclists to the pedal system. In terms of effective shoe design criteria and comfortable, injury free varticivation, an understanding of the dynamic features of these forces is important.
This rider is pedaling at 100 rpm at a power output of about 200 watts. In this figure the radiating dotted lines represent the position of the crank at one degree intervals throughout one complete revolution, that is from one TDC to the next. Positioned at the end of the crank is the pedal, the short bold line. The angle of the pedal with respect to a vertical line is called the ankling angle. The resultant forces that are applied to the pedals are shown by the bold arrows in the figure. The length of the arrow is proportional to the magnitude of the force, and its orientation shows the angle at which the force is applied.
To make the crank move around a circle it is necessary to apply forces that have a component that is perpendicular to the crank. Lafortune and Cavanaghl have termed this the effective component. The product of the effective force and the crank arm length is called torque. It is positive torque that makes the wheel go around. In Figure la the forces are labelled Fr for the total resultant force, Fn as the component of the resultant that is applied normal to the crank (the effective component), and Ft as the component that is applied parallel to the crank.
Figure 1a Effective Force
Schematic drawing of the resultant pedal force applied to the pedal during steady-rate riding. In the resultant force (Fr), the effective force (Fn), and the shear force (Ft) with respect to the crank are shown.
From Figure 1 it is evident that the pedal forces vary continuously both in magnitude and orientation throughout the pedal stroke. The consequence of the changing orientation is a change in the proportion of the total force that is effective. For example, at the bottom of the stroke the total force is quite large but is applied almost parallel to the crank arm, where as about 90 degrees after TDC the force is applied almost perpendicular to the crank. In these examples even if the total forces were the same, the effective component would be much larger at a crank angle of 90 degrees than at bottom dead center. It may well be said that when the force is not perpendicular to the crank, it is wasted, at least in terms of propulsion. Of course, that begs the question of whether it would be anatomically or otherwise possible to apply forces such that there would always be a positive, effective component.
During the second 180 degrees of crank motion the resultant force is frequently applied opposite to the desired direction. The effective component of this force will work against the rider. Remember that both legs are moving in synchrony, but 180 degrees out of phase. That is, as the left leg is moving down in propulsion, the right leg is moving up in recovery. Even though there is a small retarding force from one leg, it is easily compensated for by the other leg. During hill climbing, sprinting or other forms of low inertia riding it is likely that these recovery forces would indicate that the riders were actually pulling up on the crank. Pedaling a bicycle has been considered to occur in two phases: a propulsion phase during which the forces are applied to the first 180 degrees of crank rotation, and the recovery phase, usually considered to be the second 180 degrees. The presentation in Figure 1 tells us that propulsion occurs whenever the effective component of the applied force is positive and in the direction of pedaling. In Figure 1 it is clear that the resultant force is effective beyond the 180 degrees after TDC. Indeed, the desired goal in cycling is to make the propulsive phase as long as possible. It is instructive to note that during normal steady rate cycling, even at a Tour de France time trial pace, the pedal forces will be less than body weight as evidenced by the fact that the rider remains seated. If these forces were greater than bodyweight, he would be lifted off the seat unless he used his arms and legs to hold himself down. This has important consequences in the design of cycling footwear.
The bicycle rider applies propulsive forces to the bicycle primarily through the musculature of the legs and ultirnately through the foot. The interface between the rider's foot and the bicycle pedal is the shoe. This device has to be designed to ensure comfort, safety, and effective transmission of force. There have been many recent developments in cycling footwear and, equally importantly, in the attachment mechanism between the shoe and the pedal. These developments have been geared to enhancing the comfort of the shoe and the effectiveness of the cleating mechanism. Cycling footwear can be classified under two broad categories: those with hard soles and those with soft soles.
Some manufacturers have even produced cycling footwear that contains characteristics from both of these categories. At one end of the spectrum is the competitive cyclist who uses a hard-soled shoe. The midsole of these shoes is typically made from hard wood and, more recently, plastic. It has been suggested that the hard sole facilitates the application of force to the pedal while distributing this force over a larger area of the foot while maintaining comfort and cycling performance. However, frequent application of high pressures has been associated with localized parasthesia in long-distance cycling.
This condition may be exacerbated with the traditional cage pedal with one or more straps over the foot firmly holding it to the pedal. It seems possible that the combination of a traditional hardsoled shoe and strapped pedal may in fact lead to injury and the consequent reduction in cycling performance.
At the other end of the spectrum is the running shoe, with its thick, compliant midsole. These shoes are capable of withstanding forces that are more than twice body weight during normal running. It would seem that because pedal forces are less than body weight, these shoes would be more than capable of providing a safe, comfortable effective cycling shoe. However, there are some drawbacks to these shoes that arise from their very nature. A compliant shoe may reduce the effectiveness of the force application and create a "sloppy" fit between cyclist and bike. Such a condition would certainly affect a competitive cyclist's selection of shoe type. While there has been an increasing amount of information on pressure distribution in running and walking shoes, little published research in the area of pressure distribution in cycling shoes is evident. One such investigation was by Sanderson and Cavanagh. They investigated the inshoe pressure distribution during steady rate cycling at a high intensity of about 400 watts. Their riders used a hard-soled shoe on one occasion and a softsoled running shoe on another. Using a specially designed insole with 256 discrete force measuring elements and computer acquisition software, they were able to record the variations in the pressure distribution throughout the pedaling cycle. A sample graphic three-dimensional representation of the pressure distribution in a shoe is shown in Figure 2.
Figure 2 Sample Peak Pressure in a Shoe
A three-dimensional (3D) representation of the distribution of the peak pressures over the plantar surface of the foot during normal cycling. The height of the mountains is proportional to the pressure applied.
Sanderson and Cavanagh observed that the pressure distribution was localized in the forefoot region. It was not surprising to observe that the major pressures were recorded in the forefoot, as this is the only portion of the foot in contact with the pedal. Specifically, the first metatarsal head region and the hallux were very important load bearing areas, as indicated by the high pressures. These peaks were further accentuated by the valley between the metatarsal head region and the hallux. This is shown in Figure 2 where the major peaks in the pressure distribution are localized over the hallux and the first metatarsal head.
These authors reported that the peak pressures in the fore foot region were significantly higher for the cycling shoe than for the running shoe. They also reported that there was a more subtle and potentially more important effect noted in the comparisons between the two shoe types. The distribution of pressure in the forefoot regions was more evenly distributed in the running shoe than in the cycling shoe. That is, the large peaks of pressure over the metatarsal and hallux regions were slightly attenuated, while there was a rearward expansion of the pressure distribution in the running shoe. It seemed that the midsole of the running shoe was deforming and distributing the pressure over a broader area. The running shoe did not alter the regions of importance in terms of loading but rather made small but potentially important changes in the distribution among the regions.
These data showed that, during steady speed cycling, not all regions of the foot were equally important in their interaction with the shoe. Regardless of shoe type, the most important regions were the first metatarsal head, the lesser metatarsal heads and the hallux. In the running shoe there tended to be a more even distribution of forefoot pressures as well as a small increase in midfoot pressures. Thus, popular claims that the cycling shoe results in dramatic increases in the distribution of pressure are incorrect.
The data presented by these authors suggest that the traditional hard-soled cycling shoe could result in an increased probability of localized parasthesia because of the higher pressures. If shoe/pedal design changes are to be made with the goal of decreasing the peak pressures experienced by the rider's feet, then the contact area beneath the three important regions should be maximized.
Another interesting development which meshes quite nicely with these data is the move toward a cleatless shoe. The new clip-on pedal systems remove the old straps that contributed to the increased pressure on the foot structures. The early designs have recently been modified to incorporate a strap built into the shoe to ensure that during cycling, where upward directed forces are applied to the pedal, the integrity of the shoe is maintained. This development should go a long way to making cycling shoes considerably more comfortable and therefore more effective in terms of force transmission.
With the increase in participation in cycling there has been a concomitant increase in overuse injuries and an increase in the evaluation of the cycle as a rehabilitation tool. Bicycle riding can result in large forces applied to the legs of the rider. Forces equal to body weight have been reported for steady-rate riding In the final stages of a track race cyclists have been recorded to be pedaling in excess of 100 rpm. These conditions can contribute to a preexisting problem or can develop a new problem. There are some similiarities between the lower limb mechanics of the cyclist and the runner. In Figure 4 the stance phase of a runner and the mid propulsion point of a cyclist are drawn. In both cases the right foot is bearing relatively large loads.
Figure 4: Tibia Rotation during Running and Cycling.
Internal torsions arise during the weight bearing phase of running and the propulsion phase of cycling.
Consequently, the foot structures are responding with a pronatory action at the subtalar joint with the concomitant internal rotation of the tibia. These actions lead to an internal torsion, shown by the arrow, of the tibia. While the absolute forces may be different, these mechanical actions can lead to injuries of the knee joint in both sporting actions. There is relatively little published information on the frontal plane forces and moments acting about the knee joint during steady-rate cycling. Ericson, et al. have published data showing that the normal frontal plane knee joint moment is in varus. They further showed that the magnitude of this moment, while being dependent upon the magnitude and orientation of the vertical and mediallateral pedal reaction forces, is also dependent upon the position of the knee with respect to the pedal. Recently, Francis has shown some data indicating that wedging the foot or rotating its position on the pedal can affect the position and path of motion of the knee joint. He has hypothesized that these changes could lead to a more healthy loading pattern of the tissues within and around the knee joint. Whether the claims of force reduction under a specific situation, for example in the cruciate ligaments, are valid remains to be verified by the techniques discussed above.
Cycling has been used as a common tool for diagnosis and/or rehabilitation of patients with ischemic heart disease, postoperative training after surgery to the joints of the lower limb, exercise for patients with joint disorders such as arthritis, and as an altemative to running. All of these protocols are based on the balance between work intensity and low joint forces that can be achieved in bicycle ergometry. With the exception of Ericson et al. there has been no work examining this balance. These authors showed that the compressive forces achieved in the talocrural joint were lower in cycling than walking. Further, the estimated forces in the Achilles tendon were lower in cycling than walking. They concluded that cycling was the most appropriate exercise regime for the above discussed conditions. They also showed that these forces were independent of body weight, thus suggesting that cycling is indeed an appropriate exercise for people with weight problems.
Many issues in cycling remain to be explored. One of the most interesting aspects of development is the range of pedal-cleat systems now available. It seems that straps and cleats are rapidly becoming antiques. These have been replaced by clip-on pedals, some of which hold the foot firmly on the pedal, while others allow for some movement of the foot with respect to the pedal. As yet, there is little research data on whether these systems make any significant contribution to cycling mechanics. The supposed advantage of the movable foot system is that it will reduce torsional forces in the legs during the pedal stroke as the foot pronates and supinates. While this seems a fruitful proposal, futher work will likely reveal whether the hypothesized reductions in torque are significant.