Millions of people are involved in running and jogging activities. From those, between 37–56% are injured during the period of 1 year. 1–5 Previous injuries, excessive mileage, excessive impact forces, and excessive pronation have been proposed as major reasons for the development of running injuries. 5–9,11 The sport shoe and sport surface were assumed to influence impact forces and foot pronation. Consequently, the concepts of “cushioning” and “movement (or rearfoot) control” were developed, and strategies were studied to reduce potentially harmful impact forces and foot pronation through appropriate running shoe, shoe insert, and sport surface designs. However, results of recent studies challenged the proposed association between impact forces, foot pronation, and running injuries.
Thus, the purposes of this article are:
- To critically discuss the potential association between impact forces and foot pronation and the development of running-related injuries
- To synthesize the current knowledge about actual effects of impact forces
- To synthesize the current understanding of factors associated with foot and leg movement during running
- To propose new paradigms for the understanding of the effect of impact forces and movement control during heel–toe running.
Impact forces in heel–toe running (Figure 1) are forces resulting from the collision of the heel with the ground, reaching their maximum (the impact peak) earlier than 50 ms after first contact. 10,12 The association of impact forces with musculoskeletal injuries were typically either circumstantial in nature 7 or derived from experiments using animal models. 13 However, many results of running impact-related research studies were unexpected and did not support the concept of impact forces as one prime reason for the onset of running injuries.
Unexpected results were found in biomechanical studies as summarized in Table 1. For instance, it was found that external impact force peaks were not 14,15 or only minimally 16 influenced by changes in the hardness of running shoe midsoles. Similarly, internal impact force peaks in structures of the lower extremities were only minimally and not systematically influenced by running shoe midsoles. 17,18 However, the magnitude of vertical impact force peaks varied substantially (up to 100%) for different running velocities. 19,20 Results from model calculations showed that the magnitude of joint contact forces in the lower extremities are substantially smaller (three to five times smaller) during the impact than during the active phase of running. 18,21–23 Based on these model calculations, it was speculated that normal impact forces occurring during physical activities such as running might not be a major factor in the development of injuries in running. 23
Unexpected epidemiological results indicated that runners did not show a higher incidence of osteoarthritis than nonrunners. 24–27 Running on hard surfaces did not result in an increase of running injuries if compared with running on softer surfaces. 5 Results of a prospective study 12 did not show a significant difference of short-term running injuries between subjects with high-, medium-, and low-impact force peaks (Figure 2). Surprisingly, subjects with a high loading rate in the vertical ground reaction force had significantly fewer running-related injuries than subjects with a low loading rate. 12,28
Furthermore, in certain cases, selected shock-absorbing insoles reduced the general frequency of injuries. 29,30 However, shock-absorbing insoles were not effective in reducing the incidence of stress fractures in military recruits for which they were originally designed. 29,31 The use of a viscoelastic heel pad was proposed to be effective in reducing the symptoms of Achilles tendinitis in an athletic population, 32 however, epidemiological evidence for the effectiveness of this strategy has not been provided. These unsystematic epidemiological results for the effect of shock-absorbing insoles suggest that shock absorption might not be the primary reason for the development of these studied injuries.
Biological reactions to impact loading have been studied extensively for cartilage and bone. Impact stimuli have been shown to improve bone integrity. The increase in bone mass could be explained to 68 to 81% by the loading rate applied. 33 Impact activities such as gymnastics, basketball, running, or dancing typically produce an increase in skeletal mass, while athletes involved in low impact activities such as swimming often have a low bone density. 34,35,37 Results from measurements with young female gymnasts showed an increased vertebral bone integrity compared with a moderately active control group. 36 Premenopausal women exposed over a 2-year period to an intensive impact exercise protocol showed a 4% increase in bone mineral density compared with a 1.5% decrease for the nonimpact control group. 38 Results from in vivo experiments showed that the controlled repeated application of a force with a 1 Hz signal frequency could not maintain bone mass over an 8-week period while the same procedure with a 15 Hz signal frequency stimulated substantial new bone formation. 39 If the vertical ground reaction impact forces are simplified as sinusoidal waveforms, their frequency content is generally between 10–20 Hz. These results suggest, therefore, that impact loading has, in general, a positive effect on the development and maintenance of bone, stimulating a greater osteogenic response than nonimpact loading.
The results for cartilage are more difficult to synthesize. Some studies showed negative effects of “impact loading” on cartilage. However, in some cases, the forces applied were active instead of impact forces. Sometimes, the stresses applied were much higher than those experienced in an athlete's knee during running, and/or used loading regimens were often rather severe. 13,40–43 Furthermore, assuming that cartilage adapts to the stresses acting on it, it would be reasonable to expect different effects of this impulsive loading regimen in the rabbit, a fairly sedentary animal, than in humans, who have had many years to adapt to running-related loading. 44 Experimental results from studies with no methodological concern by the author show nonconsistent results for cartilage. High loading rates showed an increase of cartilage damage in impact experiments, compared with low loading rates. 43,45 However, high performance female gymnasts showed an increase in strength of the intervertebral discs, compared with a paralleled group of moderately active females. 36 Thus the results of these studies are not conclusive when explaining the possible effects of impact loading on cartilage.
Based on these research results, one cannot conclude that impact forces are an important factor in the development of chronic and/or acute running-related injuries. Excessive impact forces may produce damage to the human musculoskeletal system. There is a window of loading in which biologic tissues react positively to the applied impact loads. Based on the current knowledge, it is speculated by the author that impact loading for bone, cartilage, and soft tissue structures falls within the acceptable window for moderate and intensive running. It is further speculated that impact loading for bone may sometimes fall outside the acceptable window for intensive running with too-short recovery periods. However, the knowledge based on these speculations is limited.
The summary of biomechanical and epidemiological results indicates that the concept of “impact forces” as a major source for running injuries is not well understood, and that the paradigm of “cushioning” to reduce the frequency or type of running injuries should be reconsidered.
Further Biomechanical Considerations
The experimental results indicate that each runner adapts to changes in shoes or surfaces. The impact forces acting on the human foot and leg can be influenced by changing the foot and leg geometry, ankle and knee joint stiffness, and/or the coupling between the soft and rigid structures of the leg, 46,47 The response of the locomotor system may well be a combination of the three strategies. The strategies of foot/leg geometry and joint stiffness have been discussed earlier. 48 However, the strategy of changing the coupling between the soft and rigid structures of the runners leg, the strategy of muscle tuning, has only recently been proposed. 12
The muscle tuning concept suggests that the impact forces during heel strike should be considered as an input signal, characterized by amplitude and frequency. This impact force signal could produce bone vibrations at high frequencies and soft tissue vibrations of the human leg (e.g., triceps surae, quadriceps, or hamstrings muscles) at frequencies that might concur with the frequencies of the input signal. Thus soft tissue vibrations are of particular interest in this context because resonance effects could occur. Such soft tissue vibrations would cost energy, would not be comfortable, and are typically very short and heavily damped during running for muscular soft tissues. Thus, it is proposed that the muscles attempt to avoid vibrations of the soft tissue by using a tuning strategy. The concept proposes that muscles would be preactivated to create a damped vibrating system. If this assumption were correct, one would expect a change in the electromyographic (EMG) signal.
Results from pilot studies in our laboratory seem to support this speculation. 12,49 Muscle activity was quantified for selected muscles using EMG sensors. Three subjects were asked to run on three surfaces with distinctly different hardness: concrete, a normal synthetic track surface, and a soft synthetic warm-up surface. The measured EMG signal was analyzed with respect to its power spectral density and its frequency content. The results of this pilot study showed a systematic change in the power spectral density and in the median frequency for all three subjects (Figure 3).
Since this preactivation requires energy, one would expect that different shoe sole or surface material properties would require different amounts of work when performing a specific task such as running. Such a result has been theoretically predicted with an arbitrarily defined biomechanical model in which the mechanical properties of the shoe–surface interface were systematically varied. 50 The results for one set of system characteristics (Figure 4) suggested that soft and viscous materials require less work than hard and elastic ones for this specific set of characteristics.
Furthermore, one would expect that the required energy in a specific shoe situation would be subject specific since the necessary vibration damping would depend on the characteristics of the input signal and on the vibration characteristics of each soft “tissue package” (e.g., hamstrings) of a subject. Pilot results from initial experimental work support this line of thought. Ten subjects were exposed to treadmill running with VO2 quantification (measurements on four different days) in two different shoe conditions. The shoes were identical except in the material of the heel. One shoe had an elastic heel material, the other a viscous heel material. The results (Figure 5) indicate that the required work is subject and shoe specific and that the differences could be as high as 5%. Thus, the different “feeling” while running on soft or hard surfaces or shoes may be associated with changes in muscle activities and related changes in the soft tissue vibrations, both influenced by the impact input.
A Proposed New Concept/Paradigm for Impact Forces: Muscle Tuning
Based on the above considerations, a new concept/paradigm for impact forces during running is proposed:
- Impact forces are an input signal into the human body.
- This signal produces a reaction of the muscles (muscle tuning).
- The muscle tuning occurs shortly before the next contact with the ground.
- The cost function for this muscle reaction is to minimize soft tissue vibration.
- Based on the input signal and the subject-specific characteristics, muscle tuning can be low or high.
- Thus muscle tuning might affects fatigue, comfort, work, and performance.
Thus, impact forces during normal physical activity are not important because of potential injuries but rather because they affect fatigue, comfort, work and performance. Thus, it is proposed that muscle tuning is the dominant response of the human locomotor system to impact loading. The theoretical and experimental evidence for the proposed model is certainly not sufficient. However, further research should be used to solidify or reject the proposed new paradigm.
Running-related injuries have often been associated with the static or dynamic malalignment of the skeleton. Excessive foot varus or valgus positions have been speculated to create high loading situations and for repeated loading cycles overuse injuries. Additionally, excessive foot eversion and/or tibial rotation movements have been proposed to increase the chance of overuse syndromes such as patellofemoral pain syndromes, shin splints, Achilles tendinitis, plantar fasciitis, and stress fractures. 2,6,7 Thus the proper alignment of the skeleton has been proposed as being one of the most important functions of running shoes, shoe inserts, and orthotics. It has been proposed that overuse injuries due to excessive foot and leg movement, specifically due to excessive foot eversion, could be reduced with special shoes, shoe inserts, or orthotics by correcting, aligning, or limiting the skeletal movement of foot and leg. The postulated effects of such interventions were documented in clinical studies with the treatment and rehabilitation as the variables of interest, and in biomechanical studies with the changes in foot and leg movement as the variables of interest. The results of these studies are critically discussed and summarized in the next few paragraphs.
The support of the medial foot arch has often been proposed as one of the most important correction strategies for foot eversion/pronation. Some studies determined the effect of the positioning of such a medial arch support. Clinically two strategies were proposed, 51 a support under the arch of the foot (anterior support) and a support under the sustentaculum tali (posterior support). A medial arch support positioned from posterior to anterior (Figure 6) showed a mean reduction of about 4–5° for the initial shoe/leg eversion (eversion velocity) if compared with the condition with no medial support. However, no significant changes between medial support and no support were found for the total foot eversion. 52,53
Soft, semirigid, and rigid orthotics built to reduce foot/leg eversion and/or tibial rotation were studied by various groups. Movement corrections produced by such interventions were maximally between 2–4°. 54–58 The results of these studies showed generally small and nonconsistent reductions in foot and leg movement amplitudes for the tested interventions. Nevertheless these results were used to imply that shoe inserts or orthotics change (align) the skeletal movement or position.
However there are several experimental results, which suggest that the concept of “aligning the skeleton” with shoes and shoe inserts/orthotics should be reconsidered. First, in an epidemiological study, the lower extremity alignment was suggested not to be a major risk factor for running injuries based on alignment measurements 59 on a group of runners enrolling in a marathon training program. Second, results from a prospective study in our laboratory, with 131 runners and an average running distance of 30 km per week, showed that foot and ankle joint alignment and movement did not act as a predictor for an increase in running injuries over a 6-month period (Figure 7).
Third, results from a study using 12 subjects and 5 inserts with identical shape but different materials showed typically small, nonsystematic changes in foot and leg alignment and movement. 60 Inserts produced for some subject–insert combinations a reduction in foot and leg movement, for others a reduction in foot movement and an increase in leg movement, and for a third group an increase in foot movement and a decrease in leg movement. The results of this study indicate that use of an insert/orthotic is subject specific. Additionally they show that the functioning of a “good” insert/orthotic is not well understood.
Fourth, results from studies with bone pins in the calcaneus, the tibia, and the femur showed only small, nonsystematic effects of shoes or inserts on the kinematics of these bones during running. 63 Even more surprising, the differences in the skeletal movement between barefoot, shoes, and shoes with inserts were small and nonsystematic. The results of this study suggest that the locomotor system does not react to interventions with shoes, inserts, or orthotics by changing the skeletal movement pattern. These experimental results do not provide any evidence for the claim that shoes, inserts, or orthotics align the skeleton. Based on the results of the bone pin study for shoes and shoe inserts 63 one may even challenge the idea that a major function of shoes, shoe inserts, or orthotics consists in aligning the skeleton.
The previous paragraphs provided evidence contrary to the traditional concept that shoes, inserts, and orthotics align the skeleton. The skeleton seems to change its path of movement for a given task only minimally when exposed to an intervention (shoe, insert, or orthotic). One could argue (and support this argument with evidence) that—for a given task—the locomotor system chooses a strategy to keep the skeletal movement in a constant path. A similar “minimal resistance movement path” has been discussed earlier for joint movement. 62 The neuromuscular system is programed to avoid any deviation from this path. Thus, appropriate muscles will be activated if any intervention tries to produce a different skeletal movement. An optimal shoe, insert, or orthotic would minimize additional (not task related) muscle work. Consequently, shoes, inserts, and orthotics would affect muscle work, which should affect fatigue, comfort, and work/performance. Thus if comfort is an indicator for muscle activity, wearing comfortable shoes should require lower oxygen consumption than wearing uncomfortable shoes. Pilot evidence for these effects has been quantified. For instance, oxygen consumption measurements for 10 subjects running in a most and a least comfortable shoe (chosen out of five shoes provided) showed a significant difference, with more oxygen needed for running in the least comfortable shoe (Figure 8).
Thus, the behavior of a subject in a given situation (movement task and footwear) is determined by a set of factors: A force signal acts as an input variable on the shoe.
- The shoe sole acts as a first filter for the force input signal.
- The insert or orthotic acts as a second filter for the force input signal.
- The plantar surface of the foot with its mechanoreceptors senses the force–input signal.
- The signal information is transferred to the central nervous system, which provides a dynamic response based on the subject-specific conditions.
- The subject performs the movement for the task at hand.
The first three steps are situation dependent and can be influenced by the selection of the movement task, the shoe, and the insert or orthotic. The last three steps are subject dependent. The sensitivity for the mechanical signals, the potentially wobbling soft tissue masses, and the cost functions for the movement selection may be different for each subject–shoe-insert condition. This line of reasoning could explain the small and nonconsistent changes in foot and tibia movement between barefoot, shoes, and shoes with inserts 63 and the highly subject-specific differences in oxygen consumption when running with viscous or elastic heels (Figure 5). However, the experimental results and the theoretical consideration illustrate that the knowledge of the subject-specific characteristics and the appropriate insert strategies are important pieces of information for a podiatrist or orthotist when preparing an insert or orthosis for an athlete or a patient.
A Proposed New Concept/Paradigm for Foot Pronation and Movement Control
Based on current state-of-the-art knowledge, a new concept/paradigm for foot pronation is proposed:
- Forces acting on the foot during the stance phase act as an input signal.
- The locomotor system reacts to these forces by adapting the muscle activity.
- The cost function used in this adaptation process is to maintain a preferred joint movement path for a given movement task (e.g., running).
- If an intervention supports the preferred movement path, muscle activity can be reduced. If an intervention counteracts the preferred movement path, muscle activity must be increased. An optimal shoe, insert, or orthotic reduces muscle activity.
- Thus shoes, inserts, and orthotics affect general muscle activity and, therefore, fatigue, comfort, work, and performance.
Thus, “movement control” during the stance phase is not important to align the skeleton but rather because strategies to control movement change muscle activity during the stance phase. This change in muscle activity (that is not related to the actual movement task) might affect fatigue, comfort, work, and performance.
However, the presented concept/paradigm for the function of shoes, inserts, and orthotics needs more evidence and stronger evidence to support or reject it.
The two aspects of running biomechanics, “impact” and “movement control,” have been discussed, and a new way of thinking about them has been introduced. The proposed solution suggests that the locomotor system use a similar strategy in both situations, “impact” and “movement control”. In both cases the locomotor system keeps the general kinematic and kinetic situation similar for a given task. To deal with impact forces, the muscles are pre-tuned to possibly minimize soft tissue vibrations. This strategy affects the muscle activation before ground contact. To deal with shoes, inserts, and orthotics, the muscles are activated (if necessary) to provide a constant joint movement pattern. This strategy affects muscle activation during ground contact.
The characteristics of individual subjects with respect to resonance frequencies of soft tissue packages and preferred joint movement paths are different. Thus, subject-specific reactions to shoes, inserts, and orthotics are experimentally measured. However, there is some initial evidence that the signal–response pattern is similar for groups of subjects. The goal of future research should be to match subject characteristics (foot shape, lower extremity alignment, muscle strength, joint compliance, foot sensitivity, etc.) with shoe, insert, and orthotic characteristics (material properties, shape, time behavior, etc.) to find optimal group solutions for shoes, inserts, and orthotics. Based on initial results it is suggested that the needs of a large segment of the population can be served with four to five specific groups.
1. Krissoff WB, Ferris WD. Runner's injuries. Phys Sports Med 1979; 7:55–64.
2. Cavanagh PR. The Running Shoe Book.
Mountain View, CA: Anderson World, Inc., 1980.
3. Frederick EC, Hagy JL. Factors affecting peak vertical ground reaction forces in running. J Sport Biomech 1986; 2:41–49.
4. Matheson GO, Clement DB, McKenzie DC, et al. Stress fractures in athletes. A study of 320 cases. Am J Sports Med 1987; 15:46–58.
5. van Mechelen W. Running injuries: a review of the epidemiological literature. Sports Med 1992; 14:320–335.
6. Nigg BM, Eberle G, Frei D, et al. Bewegungsanalyse für Schuhkorrekturen. (Movement analysis for shoe corrections). Medita 1977; 9a:160–163.
7. James SL, Bates BT, Osternig LR. Injuries to runners. Am J Sport Med 1978; 6:40–50.
8. Cavanagh PR, Lafortune MA. Ground reaction forces in distance running. J Biomech 1980; 13:397–406.
9. Cook SD, Brinker MR, Mahlon P. Running shoes: their relation to running injuries. Sports Med 1990; 10:1–8.
10. Frederick EC, Hagy JL, Mann RA. Prediction of vertical impact force during running. J Biomech 1981; 14:498.
11. Robbins SE, Gouw GJ. Athletic footwear and chronic overloading. Sports Med 1990; 9:76–85.
12. Nigg BM. Impact forces
in running. Curr Opin Orthop 1997; 8:43–47.
13. Radin EL, Parker HG, Pugh JW, et al. Response of joints to impact loading—III. J Biomech 1973; 6:51–57.
14. Clarke TE, Frederick EC, Cooper LB. Effects of shoe cushioning upon ground reaction forces in running. Int J Sports Med 1983; 4:247–251.
15. Nigg BM, Luethi SM, Denoth J, et al. Methodological aspects of sport shoe and sport surface analysis. In:Biomechanics
. VIII-B. Matsui H, Kobayashi K, eds. Champaign, IL: Human Kinetic Publishers, 1983:1041–1052.
16. Bates BT. Comment on the influence of running velocity and midsole hardness on external impact forces
in heel-toe running. J Biomech 1989; 22:963–965.
17. Morlock M. A Generalized Three-Dimensional Six-Segment Model of the Ankle and the Foot
. PhD Thesis. The University of Calgary, Calgary, Alberta, Canada, 1990.
18. Cole GK, Nigg BM, Fick GH, et al. Internal loading of the foot and ankle during impact in running. J Appl Biomech 1995; 11:25–46.
19. Hamill J, Bates BT, Knutzen KM, et al. Variations in ground reaction force parameters at different running speeds. Hum Mov Sci 1983; 2:47–56.
20. Nigg BM, Bahlsen HA, Luethi SM, et al. The influence of running velocity and midsole hardness on external impact forces
in heel-toe running. J Biomech 1987; 20:951–959.
21. Burdett RG. Forces predicted at the ankle joint during running. Med Sci Sports Exerc 1982; 14:308–316.
22. Harrison RN, Lees A, McCullagh PJJ, et al. A bioengineering analysis of human muscle and joint forces in the lower limbs during running. J Sport Sci 1986; 4:201–218.
23. Scott SH, Winter DA. Internal forces at chronic running injury sites. Med Sci Sports Exerc 1990; 22:357–369.
24. Lane NE, Bloch DA, Jones HH, et al. Long-distance running, bone density and osteoarthritis. JAMA 1986; 255:1147–1151.
25. Panush RS, Schmidt C, Caldwell JR, et al. Is running associated with degenerative joint disease? JAMA 1986; 255:1152–1154.
26. Eichner ER. Does running cause osteoarthritis? Phys Sportsmed 1989; 17:147–154.
27. Konradsen L, Berg-Hansen EM, Söndergaard L. Long distance running and osteoarthritis. Am J Sports Med 1990; 18:379–381.
28. Bahlsen A. The Etiology of Running Injuries: A Longitudinal, Prospective Study
. PhD Thesis. The University of Calgary, Calgary, Alberta, Canada, 1989.
29. Schwellnus MP, Jordaan G, Noakes TD. Prevention of common overuse injuries by the use of shock absorbing insoles. A prospective study. Am J Sports Med 1990; 18:636–641.
30. Milgrom C, Finestone A, Shlamkovitch N, et al. Prevention of overuse injuries of the foot by improved shoe shock attenuation. A randomized prospective study. Clin Orthop Rel Res 1992; 281:189–192.
31. Gardner LI, Dziados JE, Jones BH, et al. Prevention of lower extremity stress fractures: a controlled study of a shock absorbent insole. Am J Public Health 1988; 78:1563–1567.
32. MacLellan GE, Vyvyan B. Management of pain beneath the heel and Achilles tendonitis with visco-elastic heel inserts. Br J Sports Med 1981; 15:117–121.
33. O'Connor JA, Lanyon LE. The influence of strain rate on adaptive bone remodelling. J Biomech 1982; 15:767–781.
34. Grimston SK, Zernicke RF. Exercise related stress responses in bone. J Appl Biomech 1993; 9:2–14.
35. Gross TS. “Isolating strain parameters correlated to skeletal adaptation.” Unpublished Thesis. State University of New York at Stony Brook, Stony Brook, NY, 1993.
36. Brüggemann G-P, Krämer U, Assheuer J. Vertebral Body and Disc Adaptation Due to Mechanical Loading in Artistic Gymnastics. Proceedings: 14th Sportwissenschaftlichen Hochschultag der Deutschen Vereinigung für Sportwissenschaft.
Heildeberg, Germany: Deutsche Vereinigung für Sportwissenschaft, 1999;201.
37. Dook JE, James C, Henderson NK, et al. Exercise and bone mineral density in mature female athletes. Med Sci Sports Exerc 1997; 29:291–296.
38. Heinonen A, Kannus P, Sievanen H, et al. Randomised controlled trial of effect of high-impact exercise on selected risk factors for osteoporotic fractures. Lancet 1996; 348:1343–1347.
39. McLeod KJ, Bain SD, Rubin CT. Dependency of bone adaptation on the frequency of induced dynamic strains. Trans Orthop Res Soc
40. Radin EL, Paul IL. Response of joints to impact loading. Arthritis Rheum 1971; 14:356–362.
41. Serink MT, Nachemson A, Hansson J. Effect of impact loading on rabbit knee joints. Acta Orthop Scand 1977; 48:250–262.
42. Dekel S, Weissman SL. Joint changes after overuse and peak overloading of rabbit knees in vivo. Acta Orthop Scand 1978; 49:519–528.
43. Anderson DD, Brown TD, Yang KH, et al. A dynamic finite element analysis of impulsive loading of the extension-splinted rabbit knee. J Biomech Eng 1990; 112:119–128.
44. Nigg BM, Cole GK, Brüggemann GP. Impact forces
during heel-toe running. J Appl Biomech 1995; 11:407–432.
45. Yang KH, Boyd RD, Kish VL, et al. Differential effect of load magnitude and rate on the initiation and progression of osteoarthrosis. Trans Orthop Res Soc 1989; 14:148.
46. Denoth J. Ein mechanisches Modell zur Beschreibung von passiven Belastungen (A mechanical model for the description of impact loading). In: Nigg BM, Denoth J. Sportplatzbeläge
(playing surfaces). Zurich: Juris Verlag, 1980;45–53.
47. Lafortune MA, Hennig EM, Lake MJ. Dominant role of interface over knee angle for cushioning impact loading and regulating initial leg stiffness. J Biomech 1996; 29:1523–1529.
48. Denoth J, Nigg BM. Results. In:Biomechanische Aspekte zu Sportplatzbelägen
(Biomechanical Aspects of Sport Surfaces). Nigg BM, ed. Zurich: Juris Verlag, 1978;46–47.
49. O'Flynn B. “Frequency analysis of anticipatory EMG activation of the gastrocnemius medialis and soleus while running on surfaces of different hardness.” Unpublished Master's Thesis. The University of Calgary, Canada, 1996.
50. Nigg BM, Anton M. Energy aspects for elastic and viscous shoe soles and playing surfaces. Med Sci Sports Exerc 1995; 27:92–97.
51. Segesser B, Ruepp R, Nigg BM. Indikation, Technik and Fehlermöglichkeit einer Sportschuhkorrektur (Indication, technique and error possibilities of a sport shoe correction). Orthopaedische Praxis 1978; 11:834–837.
52. Nigg BM, Bahlsen AH, Denoth J, et al. Factors influencing kinetic and kinematic variables in running. In: Biomechanics of Running Shoes. Nig BM, ed. Champaign, IL: Human Kinetics Publishers, 1986;139–159.
53. Nigg BM, Morlock M. The influence of lateral heel flare of running shoes on pronation and impact forces
. Med Sci Sports Exerc 1987; 19:294–302.
54. Smith LS, Clarke TE, Santopietro F, et al. The effects of soft and semi-rigid orthotics upon rearfoot movement in running. J Am Pod Med Assoc 1986; 76:227–233.
55. Gross ML, Davlin LB, Evanski PM. Effectiveness of orthotic shoe inserts in the long-distance runner. Am J Sports Med 1991; 19:409–412.
56. McCulloch M, Brunt D, Linden DV. The effect of foot orthotics and gait velocity on lower limb kinematics and temporal events of stance. JOSPT 1993; 17:2–10.
57. Eng JJ, Pierrynowski MR. The effect of soft foot orthotics on three-dimensional lower-limb kinematics during walking and running. Phys Ther 1994; 74:836–44.
58. Nawoczenski DA, Cook TM, Saltzman CL. The effect of foot orthotics on three-dimensional kinematics of the leg and rearfoot during running. JOSPT 1995; 21:317–327.
59. Wen DY, Puffer JC, Schmalzried TP. Lower extremity alignment and risk of overuse injuries in runners. Med Sci Sports Exerc 1997; 29:1291–1298.
60. Nigg BM, Kahn A, Fisher V, et al. Effect of shoe insert construction on foot and leg movement. Med Sci Sports Exerc 1998; 30:550–555.
61. Reinschmidt C, van den Bogert AJ, Murphy N, et al. Tibiocalcaneal motion during running—measured with external and bone markers. Clin Biomech 1997; 12:8–16.
62. Wilson DR, Feikes JD, Zavatsky AB, et al. The one degree-of-freedom nature of the human knee joint—basis for a kinematic model. Proceedings, Ninth Biennial Conference.
Vancouver: Canadian Society for Biomechanics, 1996:194–195.
63. Stacoff A, Reinschmidt C, Nigg BM, et al. Effects of foot orthoses on skeletal motion during running. Clin Biomech 2000; 15:54–64.