- Email: [email protected]
In vivo hip joint loads and pedal forces during ergometer cycling
Accepted Manuscript In vivo hip joint loads and pedal forces during ergometer cycling P. Damm, J. Dymke, A. Bender, G. Duda, G. Bergmann PII: DOI: Reference:
S0021-9290(17)30357-3 http://dx.doi.org/10.1016/j.jbiomech.2017.06.047 BM 8286
To appear in:
Journal of Biomechanics
Accepted Date:
25 June 2017
Please cite this article as: P. Damm, J. Dymke, A. Bender, G. Duda, G. Bergmann, In vivo hip joint loads and pedal forces during ergometer cycling, Journal of Biomechanics (2017), doi: http://dx.doi.org/10.1016/j.jbiomech. 2017.06.047
This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
In vivo hip joint loads and pedal forces during ergometer cycling
P. Damm, J. Dymke, A. Bender, G. Duda, G. Bergmann
Julius Wolff Institute Charité – Universitätsmedizin Berlin Augustenburger Platz 1 13353 Berlin Germany *
Corresponding author:
Philipp Damm (PhD) Julius Wolff Institute Charité - Universitätsmedizin Berlin Augustenburger Platz 1 13353 Berlin, Germany Phone: +49 30 450 559086 Email: [email protected]
Number of words Abstract: 211 Manuscript: 4555
Abstract The rising prevalence of osteoarthritis and an increase in total hip replacements calls for attention to potential therapeutic activities. Cycling is considered as a low impact exercise for the hip joint and hence recommended. However, there are limited data about hip joint loading to support this claim. The aim of this study was to measure synchronously the in vivo hip joint loads and pedal forces during cycling. The in vivo hip joint loads were measured in 5 patients with instrumented hip implants. Data were collected at several combinations of power and cadence, at two saddle heights. Joint loads and pedal forces showed strong linear correlation with power. So the relationship between the external pedal forces and internal joint forces was shown. While cycling at different cadences the minimum joint loads were acquired at 60 RPM. The lower saddle height configuration results in an approximately 15% increase compared to normal saddle height. The results offered new insights into the actual effects of cycling on the hip joint and can serve as useful tools while developing an optimum cycling regimen for individuals with coxarthrosis or following total hip arthroplasty. Due to the relatively low contact forces, cycling at a moderate power level of 90Watt at a normal saddle height is suitable for patients.
Introduction The hip joint is an integral part of the skeletal system and plays an important role in all daily activities. At an older age, this joint is often a victim of osteoarthritis, which is one of the ten most disabling diseases in the developed countries and its prevalence is on the rise (OECD, 2011). The reasons for this increase are the aging population and the growing occurrence of obesity, which is a significant risk factor. Currently the number of people opting for a total hip arthroplasty is on a rise indicating the need for more research in this field. More knowledge about the biomechanics of the hip joint implant during in vivo conditions can benefit on-going and future research in varying topics like the design of the hip joint implant, the ideal candidate for the surgery and most importantly the rehabilitation therapies after the surgery. This study investigates the loads acting on an artificial hip joint during cycling at different combinations of pedalling rate and power and at two different saddle heights. Cycling is the specific sport that this study is concentrating on because this activity is considered to have a low-impact level and has received approval by most professionals in the field as an allowed and even recommended sport after hip surgery (Klein et al., 2007; Swanson et al., 2009). Through this study, we measured the in vivo hip joint loads and corresponding pedal forces. These data will help us to identify trends and effects and consequently determine the optimum power setting and cadence for recovering patients. For this study, younger and more athletic patients compared to the average candidates for Total Hip Arthroplasty (THA) were selected. The present study varies in several aspects compared to past studies that report hip joint loads or biomechanics of cycling. Firstly, the hip joint loads reported in this study
are direct in vivo measurements whereas in other cases inverse dynamics is used to indirectly calculate joint loads (Bini et al., 2011; Brand et al., 1982; Neptune and Herzog, 1999). Secondly, THA patients participate in this study compared to the competitive cyclists and triathletes in other studies (Neptune and Hull, 1996; Sanderson et al., 2000). Thirdly, this study is undertaken with the objective to contribute to improvement of rehabilitative strategies and is not related to cycling performance and efficiency. The average patient opting for a THA has become younger (King, 2009; Swanson et al., 2009) and leads to a more physically active lifestyle (Crowninshield et al., 2006). This means returning to athletic activities after hip surgery is an expectation and even a priority for many candidates. The goal of this study is to analyze the acquired in vivo load data and external forces of a THA patient during cycling. The focus is on the effects of a sport especially because the profile of people receiving total hip replacements has changed over time.
Methods Instrumented implant In order to obtain data about the joint loads during cycling, an instrumented implant was used, which is capable of telemetrically transferring in vivo data. The instrumented implant and the external equipment were described elsewhere (Damm et al., 2010; Graichen et al., 2007, 1994). Having implanted this device within patients gives us access to primary information about the forces and moments acting on the hip joint implant during every movement the patient undertakes. With the instrumented implants
the joint loads acting between the head and the cup, can be measured in vivo with an accuracy of 1-2%. The femur-based coordinate system (Glaser and Bornkessel, 2002) is located in the head centre of a right sided implant, but is defined relative to the bone. If the implant is in the left leg, the forces and moments are mirrored on to the right hip to keep the data consistent. The three force components are mutually perpendicular. The Z-axis is parallel to the axis of the femur; The X-axis is in the lateral-medial direction and the Yaxis in the anterior-posterior direction. The positive force components act in lateral, anterior, and superior directions. The moments act around the axes in a clockwise direction. The resultant contact force Fres is calculated by the three force components and the three moment components deliver the resultant friction moment Mres. However, only the resultant load components F res were reported here. The forces are normalized as a percentage of bodyweight (% BW) for better inter-patient comparability. Instrumented pedals For the synchronous measurement of the pedal forces the ‘Powerforce’ system (Radlabor, Freiburg, Germany) was used (Stapelfeldt et al., 2007). It is mounted between the crank and pedal and measure forces separately for the left and right foot. The system was connected to the external equipment by cables. At each pedal, there are two sensor systems which are based on the Hall-Effect. A sensor and a magnet are used; the sensor measures the magnetic field variations as it displaces in respect to the magnet. The two systems are orthogonally located, allowing the decoupled measurement of a radial and tangential force. The radial force acts parallel to the crank, causing very slight lengthening and compressing of the crank,
rendering it useless, hence called the loss force. The tangential force acts perpendicular to the crank, contributing to the forward thrust movement of the crank in the pedalling direction; therefore it is also called the effective force. However, only the resultant pedal forces at the ipsilateral side (F pi) and at the contralateral side (F pc) were reported here. The patients are videotaped during the whole measurement. The telemetric load signals and the pedal forces are recorded simultaneously. Subjects and measurements Five male subjects with instrumented hip joint prostheses participated at this study (Table 1). They gave their informed written consent to participate. The study was approved by the ethical committee of the Charité – Universitätsmedizin Berlin, Germany (EA2/057/09)
and
was
registered
in
the
German
Clinical
Trials
Register
(DRKS00000563). They performed during measurement cycling regimen with different cadences, powers and saddle heights (Table 2). The ‘normal’ saddle height was defined with an individual knee flexion angle between 30-40° at the lowest point of the pedal crank. For the 'low' position, the height of the saddle was decreased by 9cm. The data were acquired during measurement sessions where each subject performed cycling on a pre-programmed ergometer. Every ergometer setup (power level, cadence, seat height) was performed for 2 minutes. The repetitive action during cycling is the extension of joints followed by flexion and the corresponding revolution of the crank arm and pedal. One complete rotation corresponds to a single load cycle. The resultant in vivo force acting on the hip joint and the pedal force during this are shown in figure 1. The beginning of every load cycle is when the crank arm is vertical (0°) and the pedal is at its highest position. The first half
of the rotation is the power phase when the force for propulsion is generated. The second half is known as the recovery phase. The position of the crank arm as a function of the joint force were taken from the individual videos. Data evaluation All forces are reported in percent of the bodyweight (% BW). Average force-time patterns from all measured load cycles were calculated for each subject. The employed ‘time warping’ method (Bender and Bergmann, 2011) first normalises the period times of all included load cycles. Then the single time scales are distorted in such a way that the squared differences between all deformed curves, summed over the whole cycle time, become a minimum. Finally, an arithmetically averaged load-time pattern is calculated from all deformed curves. Using these algorithms, an average course was first calculated from the time patterns of F res. The obtained time deformations of the single trials were then transferred to the belonging force components before averaging them. This procedure was first applied on all trials of the single patients, leading to ‘intraindividual’ averages. Then these averages from all patients were combined to an ‘interindividual’ average, which describes the loads acting in an ‘average’ patient. The individual peak values were not taken from the averaged time courses but instead the numerical peak values of the single trials were averaged, first intra- and then interindividually. Extreme values of the averaged load-time patterns can slightly deviate from these numerically averaged numbers. The intra-individual determined numerical differences were statistical analysed (Mann-Whitney, p ≤ 0.05).
Results
Load pattern In figure 1 Fres, Fpi and Fpc are plotted for a single load cycle. F res comprises mostly of force acting in the -z direction. This force acts downwards, parallel to the femur, giving rise to compressive force in the hip joint. The maximum of F res with 99% BW on average occur when the crank arm is approximately at 90° and the minimum of 40% BW, when the crank is shortly after 180°. The peak of F pi (32% BW) coincides with the peak of the contact force; both occur around 25% of the load cycle and thus are clearly connected and have the same origin. As F pi decreases, the contralateral pedal force (Fpc) starts rising; even though the motion looks symmetrical, the pedal forces show distinct differences. In the figure 2 (A), the box plots for the maximum and minimum contact forces for each patient are shown. These plots include all the values that the cyclist reached during the 2 minutes of cycling at 40 revolutions per minute (RPM) and 90W. A wide range of peak values were determined, both inter- and intra-individually. At 40 RPM, the maximum of Fres varies from 69% BW (H9L) to 121% BW (H2R). The minimum following the peak occurs approximately half way through the load cycle and is of similar magnitudes in all patients, averaging at 40% BW. The variation in each pedal force can be seen in figure 2 (B) and the comparison between the two pedal forces can be assessed. A statistical analysis (Mann-Whitney) was performed on each patient’s data to study the intra-individual differences in the pedal forces. The differences between the ipsilateral and contralateral pedal forces were found to be statistically significant in all patients except patient H8L. However, there are inter-individual differences in the direction of this difference. Patients H6R, H7R and H9L have higher ipsilateral pedal forces and patient H2R has lower ipsilateral pedal forces.
Influence of power The participating patients cycled at constant cadence and increasing power levels; the speed of 40 RPM was maintained and the power increased from 50 to 130W in 20W increments. The profile of each load cycle resembles the generic pattern previously shown. The comparison of the in vivo load pattern of the average subject at the different powers curves clearly shows the increase in magnitude (figure 3A) during the first half of the cycle, the power phase. The peaks of these curves range from 67% BW at 50W power setting to 137% BW at 130W; approximately a 20% BW increase with 20W increase (Table 3). A distinct linearly increasing trend was visible for every patient. The differences between each power setting for every patient were statistically significant with a p-value of less than 0.01. Also for the average patient a high correlation (Regression Equation: y = 0.89x + 19.7) between the power level and the maximum in vivo joint force occurred (R² = 0.996) and their low p-values (p<0.01) imply a statistically significant correlation. The increase in power had also influences on the force exerted on the pedals. In figure 3B, the pedal force pattern of the average patient at every power level is plotted. The magnitude of Fpi and Fpc during the down stroke (0-50% load cycle, 50-100% load cycle respectively) is higher at higher powers. F pi ranges from 23% BW at 50W up to 41% BW at 130W. Although this is not a steep increase, it is a steady increase of about 5.5% BW in pedal forces for every power increment. The patterns of F pc at different powers are not consistent in shape during the second half of the load cycle. Additionally, they are comparatively lower than Fpi. The differences between the ipsilateral and contralateral pedal forces increased as the power increased.
Hip joint force versus pedal force A very strong correlation with power was shown for the in vivo hip joint force and the pedal forces. So a relation between the pedal forces and the in vivo forces is investigated. In figure 4 the average maximum values of the in vivo forces are plotted against pedal forces for each patient. A strong linear correlation was occurring, with a high regression coefficients in all cases (Regression Equation: y = 3.23x - 6.01). The relation between external forces and internal forces is established by this data collected at the cadence of 40 RPM and varying power levels. Influence of cadence The influence of cadence on the contact forces and pedal forces of the average patient is shown in figure 5. Lesser force is required while cycling at higher cadences. In figure 5A the load pattern of 40 RPM has the highest peak magnitude at 99% BW while the lowest peak occurs at 73% BW for cadences of 60 RPM. However, at 90 RPM, the peak magnitude does not conform to this trend and has a high peak magnitude of 93% BW. This effect was seen in all patients, except H2R. The effect of cadence on pedal forces is not parallel to the effect seen in the in vivo forces. A more distinct trend appears when pedal forces are analyzed. In figure 5B the maximum pedal forces at different cadences were plotted for the average patient. It is shown, that the peak magnitude of F pi decreases as the cadence increases, ranging from 32% BW at 40 RPM to 23% BW at 90 RPM. Influence of seat height The ergometer cycling that was performed at specific powers by the patients at a normal saddle height was repeated at a lower saddle height (-9cm). In table 3, the values of the
in vivo hip joint force and the corresponding pedal forces for the average patient at different powers and different seat heights and the percentage increase are listed. The contact forces at low saddle height are slightly increased to the forces at normal saddle height, between 7-15%. The pedal forces at low saddle height are slightly higher (1-2% BW) than the pedal forces at normal saddle height.
Discussion The goal of this study was to analyze the acquired in vivo load data and external forces of a THA patient during cycling. The focus was on the effects of a sport especially because the profile of people receiving total hip replacements has changed over time. Influence of power Ergometer cycling at different power outputs has been extensively studied and various factors have been investigated; muscle activity, muscular power output (Ericson, 1988), pedal force application and in-shoe pressure distribution (Sanderson et al., 2000). In several studies, pedal forces have shown to increase with power, which was confirmed by the present study too. At higher power settings, positive hip extensor and flexor power output increases, the former at a higher rate than the latter (Ericson, 1988). Our results confirm that these effects can be translated to hip joint loads, since contact forces increased with higher power levels. Workload has been established as the most important variable to engender change in joint loads and muscular activity (Ericson et al., 1986). Influence of cadence
The power generated on an ergometer is the product of the crank torque and angular velocity. Therefore, to generate a specific power, the combinations of the applied force and the pedalling rate are important. Physical laws state that as cadence is increased, lesser force is required to maintain a constant power output. This effect holds for the pedal forces. At a constant power output of 90 W, pedal forces linearly decreased from 32 to 23% BW as cadence increased from 40 to 90 RPM. However this trend is not reflected in the in vivo loads at higher cadences. The contact forces at the joint do reduce as the cadence shifts from 40 to 60 RPM, however at 90 RPM a deviation is encountered, where the magnitude rises. Several studies have reported changes in muscle work (Neptune and Herzog, 1999), muscle co-ordination (Li and Caldwell, 1997; Neptune and Hull, 1996) and muscle activation dynamics (Neptune and Kautz, 2001) in the lower extremity during changes in cadences. Consequently, the changes in muscle activity have been well documented in cycling literature. However this study provides an insight into the effects of pedalling speed on in vivo joint forces. Previous studies on cadence effects can provide an explanation for the effect changing cadence is having on the hip joint loads. The increased magnitude of in vivo loads at higher pedalling speeds is the consequence of higher force generation by muscles and co-contraction of muscles. Activation dynamics come into play at higher cadences because of rapid contraction-relaxation cycles that muscles need to undergo. The delay between excitation and force generation, gives rise to force during the relaxation phase (lengthening of the muscle) rendering it useless; also called negative muscle work (Neptune and Herzog, 1999). To overcome this and maintain a constant power output, the muscles have to produce more positive work, leading to a higher contact force at the joint (Neptune and Herzog, 1999;
Neptune and Hull, 1996). Another cause is co-contraction of muscles at higher cadences, required for control and stabilization of joints. The preferred pedalling rate for cyclists is approximately 90 RPM and cadence effects usually begin after 90 RPM (Neptune and Hull, 1996). All previous studies have been done with healthy inexperienced cyclists. It should be noted that in this study, all participants are THA patients. The minimum in vivo contact forces were encountered at 60 RPM, therefore this cadence is recommended for rehabilitation. Due to the age, sub-optimal fitness levels and recovering status of patients, there were limitations on collecting data at higher cadences than 90 RPM. Influence of saddle height The effect of saddle height adjustments on pedal forces has previously been reported. Lower saddle heights produced higher resultant pedal forces but lower force effectiveness (Bini et al., 2011). In this study also, slightly higher pedal forces were encountered at low saddle height configuration at most cycling settings. Studies to date have included triathletes and cyclists with competitive experiences and the effects on joint angles and joint mechanical work have been collected or derived (Bini et al., 2011). Moreover, little research has been done regarding the hip joint, with more importance given to the knee joint. This study is the first to look into the saddle height effects on in vivo hip joint loads and uses data from cyclists with THA. The contact forces arising within the hip joint are slightly higher at a low saddle height; there is approximately an increase in contact forces between 7-15% on average. A study of muscle recruitment pattern during different saddle heights showed an increase in muscle activity with lower saddle heights. This trend was more prevalent in quadriceps and hamstrings (Raymond et al., 2005). The increased muscle forces at lower saddle
height are a possible explanation for increased hip joint contact forces. For hip arthroplasty
patients,
therapeutic
ergometer
cycling,
during
rehabilitation
is
recommended to be done at a normal saddle height as lower contact forces can be expected. Conclusion The in vivo measurements during cycling with instrumented hip implants have shown that comparing with other activities, cycling is an athletic activity with low-impact to the hip joint. During cycling inter-patient differences were observed in F res, this can be attributed to several factors like post-operative time, fitness level, agility and weight. For example, patient H9L was considerably heavier than the rest of the patients, which explain the high deviation of his contact forces from the rest of the patients. Additionally, there was a high inter-individual difference in post operative time; varying from 6 to 30 months. Besides the high inter-individual differences, there exists a wide range of contact forces even intra- individually, which can be a result of change in style and/or fatigue during the cycling measurement. When comparing the forces applied on both ipsilateral and contralateral pedals, it became undeniable that the operated and normal hip joints and surrounding muscles are not used equally in any patient. These inter-individual differences can be attributed to several factors like agility and athletic activities in daily living and also the patient’s adherence to the rehabilitation program. These factors ultimately affect the recovery of musculature around the hip joint. For example, a study compared the outcomes of patients receiving weight-bearing and postural stability exercises and patients doing basic isometric and active range of motion exercises 4 months after THA. The former
had significantly improved muscle strength, postural stability and self-perceived function (Trudelle-Jackson and Smith, 2004). Another factor to consider is whether the patient’s operated leg is the same as the dominant leg. All patients except H7R, had their dominant leg operated, i.e. their ipsilateral leg is their dominant leg. In figure 2B for patient H7R, the difference between the ipsilateral and contralateral pedal forces is greater than any other patient. Limitation The present study’s main objective is to measure the actual hip joint loads in vivo, which is accomplished with patients recovering from total hip replacements. The participants in this study are comparatively have lower fitness levels and are less experienced than the typical athletic cyclist is. Due to these reasons and the fact that the hip joints are different in the groups of participants (natural and artificial), limitations exist while correlating data. The small number of subjects participating in this study serves as a limitation.
Additional Data Selected examples of the in vivo measurements, on which this study was based, are published in the public data base www.orthoload.com.
Acknowledgements This work was supported by Deutsche Forschungsgemeinschaft (Be 804/19-1), the German Federal Ministry of Education and Research (BMBF 01EC1408A; OverloadPrevOP; SPO3), Deutsche Arthrose-Hilfe and the OrthoLoadClub.
Literature
Bender, A., Bergmann, G., 2011. Determination of Typical Patterns from Strongly Varying Signals. Computer Methods in Biomechanics and Biomedical Engineering iFirst, 1–9.
Bini, R.R., Hume, P.A., Crofta, J.L., 2011. Effects of saddle height on pedal force effectiveness. Engineering Procedia 13, 51–55.
Brand, R.A., Crowninshield, R. D., Pedersen, D.R., 1982. Forces on the femoral head during acitivties of daily living. The Iova Orthopaedic Journal 2, 43–49.
Crowninshield, Roy D, Rosenberg, A.G., Sporer, S.M., 2006. Changing demographics of patients with total joint replacement. Clinical Orthopaedics and Related Research 3.
Damm, P., Graichen, Friedmar, Rohlmann, Antonius, Bender, Alwina, Bergmann, Georg, 2010. Total hip joint prosthesis for in vivo measurement of forces and moments. Medical engineering & physics 32, 95–100.
Ericson, M.O., 1988. Mechanical muscular power output and work during ergometer cycling at different work loads and speeds. European Journal of Applied Physiology and Occupational Physiology 57, 382–387.
Ericson, M.O., Bratt, A., Nisell, R., Arborelius, U.P., Ekholm, J., 1986. Applied Physiology Power output and work in different muscle groups during ergometer cycling.
Glaser, R., Bornkessel, C., 2002. ISB recommendation on definitions of joint coordinate system of various joints for the reporting of human motion - part I: ankle, hip, and spine. Journal of Biomechanics 35, 543–548.
Graichen, F., Arnold, R., Rohlmann, A., Bergmann, G., 2007. Implantable 9-channel telemetry system for in vivo load measurements with orthopedic implants. IEEE transactions on bio-medical engineering 54, 253–61.
Graichen, F., Bergmann, G., Rohlmann, A., 1994. Telemetric transmission system for in vivo measurement of the stress load of an internal spinal fixator. Biomedizinische Technik Biomedical engineering 39, 251–258.
King, A.E., 2009. Current Trends in Total Hip Replacement for Younger Patients. Journal of the Canadian Rheumatology Association 19, 26–27. Klein, G.R., Levine, B.R., Hozack, W.J., Strauss, E.J., D’Antonio, J. a, Macaulay, W., Di Cesare, P.E., 2007. Return to athletic activity after total hip arthroplasty. Consensus guidelines based on a survey of the Hip Society and American Association of Hip and Knee Surgeons. The Journal of arthroplasty 22, 171–5.
Li, L., Caldwell, G.E., 1997. The Effect of cyling cadences on the coordination of monoand bi-articular muscles. IEEE transactions on bio-medical engineering 334–337.
Neptune, R., Herzog, W., 1999. The association between negative muscle work and pedaling rate. Journal of Biomechanics 32, 1021–1026.
Neptune, R., Hull, M.L., 1996. Methods for Determining Hip Movement in Seated cycling and Their Effect on Kinematics and Kinetics 493–507.
Neptune, R., Kautz, S., 2001. Muscle activation and deactivation dynamics: the governing properties in fast cyclical human movement performance? Exercise and sport sciences reviews 29, 76–80.
OECD, 2011. Hip and knee replacement. Healt at a Glance OECD Indic.
Okoro, T., Lemmey, A.B., Maddison, P., Andrew, J.. ., 2012. An appraisal of rehabilitation regimes used for improving functional outcome after total hip replacement surgery. Sports medicine, arthroscopy, rehabilitation, therapy & technology 4.
Raymond, C.H.S., Ng, J.K.-F., Ng, G.Y.F., 2005. Muscle recruitment pattern in cycling: a review. Physical Therapy in Sport 6, 89–96.
Sanderson, D.J., Hennig, E.M., Black, A.H., 2000. The influence of cadence and power output on force application and in-shoe pressure distribution during cycling by competitive and recreational cyclists. Journal of sports sciences 18, 173–81.
Stapelfeldt, B., Mornieux, G., Oberheim, R., Belli, A., Gollhofer, A., 2007. Development and evaluation of a new bicycle instrument for measurements of pedal forces and power output in cycling. International journal of sports medicine 28, 326–32.
Swanson, E. a, Schmalzried, T.P., Dorey, F.J., 2009. Activity recommendations after total hip and knee arthroplasty: a survey of the American Association for Hip and Knee Surgeons. The Journal of arthroplasty 24, 120–6.
Trudelle-Jackson, E., Smith, S., 2004. Effects of a late-phase exercise program after total hip arthroplasty: a randomized controlled trial. Archives of Physical Medicine and Rehabilitation 85, 1056–1062.
Figure Legends
Figure 1: In vivo hip joint loads and corresponding pedal forces; 90W 40 RPM; average subject
Figure 2: In vivo hip joint force (A) and pedal force during cycling; 40 RPM; 90W
Figure 3: Pattern of the in vivo hip joint force and pedal forces at 40 RPM and different ergometer power; average subject
Figure 4: Relation between maximum hip joint forces and the maximum ipsilateral pedal forces of each patient (coloured) and the average subject (black)
Figure 5: Maximum of the in vivo joint forces (A) and pedal forces (B); 90W;
Table 1:
Subjects investigated
Patient
H2R
H6R
H7R
H8L
H9L
Age
65
69
54
56
54
Body Weight
81
83
94
88
118
pOP Month
30
12
15
10
6
Site of THA
Right
Right
Right
Left
Left
Table 2:
Parameters adopted for cycling measurements
Saddle Height
Normal
Low [- 9cm]
Power at 40 RPM
50, 70, 90, 110 and 130W
50, 70, 90 and 110W
Cadence at 90W
40, 70 and 90 RPM
40, 70 and 90RPM
Table 3:
In vivo joint force and ipsilateral pedal forces at different seat heights and ergometer power; average subject
Ergometer
Normal Saddle Height
Low Saddle Height [- 9cm]
Δ Joint Force
Power
Joint Force
Pedal Force
Joint Force
Pedal Force
[W]
[% BW] (STD)
[% BW] (STD)
[% BW] (STD)
[% BW] (STD)
50
67 (13)
23 (3)
72 (13)
25 (3)
7
70
80 (14)
28 (3)
91 (16)
29 (3)
12
90
97 (16)
32 (3)
114 (21)
34 (3)
15
110
118 (19)
38 (3)
134 (20)
39 (3)
12
[%]
S0021-9290(17)30357-3 http://dx.doi.org/10.1016/j.jbiomech.2017.06.047 BM 8286
To appear in:
Journal of Biomechanics
Accepted Date:
25 June 2017
Please cite this article as: P. Damm, J. Dymke, A. Bender, G. Duda, G. Bergmann, In vivo hip joint loads and pedal forces during ergometer cycling, Journal of Biomechanics (2017), doi: http://dx.doi.org/10.1016/j.jbiomech. 2017.06.047
This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
In vivo hip joint loads and pedal forces during ergometer cycling
P. Damm, J. Dymke, A. Bender, G. Duda, G. Bergmann
Julius Wolff Institute Charité – Universitätsmedizin Berlin Augustenburger Platz 1 13353 Berlin Germany *
Corresponding author:
Philipp Damm (PhD) Julius Wolff Institute Charité - Universitätsmedizin Berlin Augustenburger Platz 1 13353 Berlin, Germany Phone: +49 30 450 559086 Email: [email protected]
Number of words Abstract: 211 Manuscript: 4555
Abstract The rising prevalence of osteoarthritis and an increase in total hip replacements calls for attention to potential therapeutic activities. Cycling is considered as a low impact exercise for the hip joint and hence recommended. However, there are limited data about hip joint loading to support this claim. The aim of this study was to measure synchronously the in vivo hip joint loads and pedal forces during cycling. The in vivo hip joint loads were measured in 5 patients with instrumented hip implants. Data were collected at several combinations of power and cadence, at two saddle heights. Joint loads and pedal forces showed strong linear correlation with power. So the relationship between the external pedal forces and internal joint forces was shown. While cycling at different cadences the minimum joint loads were acquired at 60 RPM. The lower saddle height configuration results in an approximately 15% increase compared to normal saddle height. The results offered new insights into the actual effects of cycling on the hip joint and can serve as useful tools while developing an optimum cycling regimen for individuals with coxarthrosis or following total hip arthroplasty. Due to the relatively low contact forces, cycling at a moderate power level of 90Watt at a normal saddle height is suitable for patients.
Introduction The hip joint is an integral part of the skeletal system and plays an important role in all daily activities. At an older age, this joint is often a victim of osteoarthritis, which is one of the ten most disabling diseases in the developed countries and its prevalence is on the rise (OECD, 2011). The reasons for this increase are the aging population and the growing occurrence of obesity, which is a significant risk factor. Currently the number of people opting for a total hip arthroplasty is on a rise indicating the need for more research in this field. More knowledge about the biomechanics of the hip joint implant during in vivo conditions can benefit on-going and future research in varying topics like the design of the hip joint implant, the ideal candidate for the surgery and most importantly the rehabilitation therapies after the surgery. This study investigates the loads acting on an artificial hip joint during cycling at different combinations of pedalling rate and power and at two different saddle heights. Cycling is the specific sport that this study is concentrating on because this activity is considered to have a low-impact level and has received approval by most professionals in the field as an allowed and even recommended sport after hip surgery (Klein et al., 2007; Swanson et al., 2009). Through this study, we measured the in vivo hip joint loads and corresponding pedal forces. These data will help us to identify trends and effects and consequently determine the optimum power setting and cadence for recovering patients. For this study, younger and more athletic patients compared to the average candidates for Total Hip Arthroplasty (THA) were selected. The present study varies in several aspects compared to past studies that report hip joint loads or biomechanics of cycling. Firstly, the hip joint loads reported in this study
are direct in vivo measurements whereas in other cases inverse dynamics is used to indirectly calculate joint loads (Bini et al., 2011; Brand et al., 1982; Neptune and Herzog, 1999). Secondly, THA patients participate in this study compared to the competitive cyclists and triathletes in other studies (Neptune and Hull, 1996; Sanderson et al., 2000). Thirdly, this study is undertaken with the objective to contribute to improvement of rehabilitative strategies and is not related to cycling performance and efficiency. The average patient opting for a THA has become younger (King, 2009; Swanson et al., 2009) and leads to a more physically active lifestyle (Crowninshield et al., 2006). This means returning to athletic activities after hip surgery is an expectation and even a priority for many candidates. The goal of this study is to analyze the acquired in vivo load data and external forces of a THA patient during cycling. The focus is on the effects of a sport especially because the profile of people receiving total hip replacements has changed over time.
Methods Instrumented implant In order to obtain data about the joint loads during cycling, an instrumented implant was used, which is capable of telemetrically transferring in vivo data. The instrumented implant and the external equipment were described elsewhere (Damm et al., 2010; Graichen et al., 2007, 1994). Having implanted this device within patients gives us access to primary information about the forces and moments acting on the hip joint implant during every movement the patient undertakes. With the instrumented implants
the joint loads acting between the head and the cup, can be measured in vivo with an accuracy of 1-2%. The femur-based coordinate system (Glaser and Bornkessel, 2002) is located in the head centre of a right sided implant, but is defined relative to the bone. If the implant is in the left leg, the forces and moments are mirrored on to the right hip to keep the data consistent. The three force components are mutually perpendicular. The Z-axis is parallel to the axis of the femur; The X-axis is in the lateral-medial direction and the Yaxis in the anterior-posterior direction. The positive force components act in lateral, anterior, and superior directions. The moments act around the axes in a clockwise direction. The resultant contact force Fres is calculated by the three force components and the three moment components deliver the resultant friction moment Mres. However, only the resultant load components F res were reported here. The forces are normalized as a percentage of bodyweight (% BW) for better inter-patient comparability. Instrumented pedals For the synchronous measurement of the pedal forces the ‘Powerforce’ system (Radlabor, Freiburg, Germany) was used (Stapelfeldt et al., 2007). It is mounted between the crank and pedal and measure forces separately for the left and right foot. The system was connected to the external equipment by cables. At each pedal, there are two sensor systems which are based on the Hall-Effect. A sensor and a magnet are used; the sensor measures the magnetic field variations as it displaces in respect to the magnet. The two systems are orthogonally located, allowing the decoupled measurement of a radial and tangential force. The radial force acts parallel to the crank, causing very slight lengthening and compressing of the crank,
rendering it useless, hence called the loss force. The tangential force acts perpendicular to the crank, contributing to the forward thrust movement of the crank in the pedalling direction; therefore it is also called the effective force. However, only the resultant pedal forces at the ipsilateral side (F pi) and at the contralateral side (F pc) were reported here. The patients are videotaped during the whole measurement. The telemetric load signals and the pedal forces are recorded simultaneously. Subjects and measurements Five male subjects with instrumented hip joint prostheses participated at this study (Table 1). They gave their informed written consent to participate. The study was approved by the ethical committee of the Charité – Universitätsmedizin Berlin, Germany (EA2/057/09)
and
was
registered
in
the
German
Clinical
Trials
Register
(DRKS00000563). They performed during measurement cycling regimen with different cadences, powers and saddle heights (Table 2). The ‘normal’ saddle height was defined with an individual knee flexion angle between 30-40° at the lowest point of the pedal crank. For the 'low' position, the height of the saddle was decreased by 9cm. The data were acquired during measurement sessions where each subject performed cycling on a pre-programmed ergometer. Every ergometer setup (power level, cadence, seat height) was performed for 2 minutes. The repetitive action during cycling is the extension of joints followed by flexion and the corresponding revolution of the crank arm and pedal. One complete rotation corresponds to a single load cycle. The resultant in vivo force acting on the hip joint and the pedal force during this are shown in figure 1. The beginning of every load cycle is when the crank arm is vertical (0°) and the pedal is at its highest position. The first half
of the rotation is the power phase when the force for propulsion is generated. The second half is known as the recovery phase. The position of the crank arm as a function of the joint force were taken from the individual videos. Data evaluation All forces are reported in percent of the bodyweight (% BW). Average force-time patterns from all measured load cycles were calculated for each subject. The employed ‘time warping’ method (Bender and Bergmann, 2011) first normalises the period times of all included load cycles. Then the single time scales are distorted in such a way that the squared differences between all deformed curves, summed over the whole cycle time, become a minimum. Finally, an arithmetically averaged load-time pattern is calculated from all deformed curves. Using these algorithms, an average course was first calculated from the time patterns of F res. The obtained time deformations of the single trials were then transferred to the belonging force components before averaging them. This procedure was first applied on all trials of the single patients, leading to ‘intraindividual’ averages. Then these averages from all patients were combined to an ‘interindividual’ average, which describes the loads acting in an ‘average’ patient. The individual peak values were not taken from the averaged time courses but instead the numerical peak values of the single trials were averaged, first intra- and then interindividually. Extreme values of the averaged load-time patterns can slightly deviate from these numerically averaged numbers. The intra-individual determined numerical differences were statistical analysed (Mann-Whitney, p ≤ 0.05).
Results
Load pattern In figure 1 Fres, Fpi and Fpc are plotted for a single load cycle. F res comprises mostly of force acting in the -z direction. This force acts downwards, parallel to the femur, giving rise to compressive force in the hip joint. The maximum of F res with 99% BW on average occur when the crank arm is approximately at 90° and the minimum of 40% BW, when the crank is shortly after 180°. The peak of F pi (32% BW) coincides with the peak of the contact force; both occur around 25% of the load cycle and thus are clearly connected and have the same origin. As F pi decreases, the contralateral pedal force (Fpc) starts rising; even though the motion looks symmetrical, the pedal forces show distinct differences. In the figure 2 (A), the box plots for the maximum and minimum contact forces for each patient are shown. These plots include all the values that the cyclist reached during the 2 minutes of cycling at 40 revolutions per minute (RPM) and 90W. A wide range of peak values were determined, both inter- and intra-individually. At 40 RPM, the maximum of Fres varies from 69% BW (H9L) to 121% BW (H2R). The minimum following the peak occurs approximately half way through the load cycle and is of similar magnitudes in all patients, averaging at 40% BW. The variation in each pedal force can be seen in figure 2 (B) and the comparison between the two pedal forces can be assessed. A statistical analysis (Mann-Whitney) was performed on each patient’s data to study the intra-individual differences in the pedal forces. The differences between the ipsilateral and contralateral pedal forces were found to be statistically significant in all patients except patient H8L. However, there are inter-individual differences in the direction of this difference. Patients H6R, H7R and H9L have higher ipsilateral pedal forces and patient H2R has lower ipsilateral pedal forces.
Influence of power The participating patients cycled at constant cadence and increasing power levels; the speed of 40 RPM was maintained and the power increased from 50 to 130W in 20W increments. The profile of each load cycle resembles the generic pattern previously shown. The comparison of the in vivo load pattern of the average subject at the different powers curves clearly shows the increase in magnitude (figure 3A) during the first half of the cycle, the power phase. The peaks of these curves range from 67% BW at 50W power setting to 137% BW at 130W; approximately a 20% BW increase with 20W increase (Table 3). A distinct linearly increasing trend was visible for every patient. The differences between each power setting for every patient were statistically significant with a p-value of less than 0.01. Also for the average patient a high correlation (Regression Equation: y = 0.89x + 19.7) between the power level and the maximum in vivo joint force occurred (R² = 0.996) and their low p-values (p<0.01) imply a statistically significant correlation. The increase in power had also influences on the force exerted on the pedals. In figure 3B, the pedal force pattern of the average patient at every power level is plotted. The magnitude of Fpi and Fpc during the down stroke (0-50% load cycle, 50-100% load cycle respectively) is higher at higher powers. F pi ranges from 23% BW at 50W up to 41% BW at 130W. Although this is not a steep increase, it is a steady increase of about 5.5% BW in pedal forces for every power increment. The patterns of F pc at different powers are not consistent in shape during the second half of the load cycle. Additionally, they are comparatively lower than Fpi. The differences between the ipsilateral and contralateral pedal forces increased as the power increased.
Hip joint force versus pedal force A very strong correlation with power was shown for the in vivo hip joint force and the pedal forces. So a relation between the pedal forces and the in vivo forces is investigated. In figure 4 the average maximum values of the in vivo forces are plotted against pedal forces for each patient. A strong linear correlation was occurring, with a high regression coefficients in all cases (Regression Equation: y = 3.23x - 6.01). The relation between external forces and internal forces is established by this data collected at the cadence of 40 RPM and varying power levels. Influence of cadence The influence of cadence on the contact forces and pedal forces of the average patient is shown in figure 5. Lesser force is required while cycling at higher cadences. In figure 5A the load pattern of 40 RPM has the highest peak magnitude at 99% BW while the lowest peak occurs at 73% BW for cadences of 60 RPM. However, at 90 RPM, the peak magnitude does not conform to this trend and has a high peak magnitude of 93% BW. This effect was seen in all patients, except H2R. The effect of cadence on pedal forces is not parallel to the effect seen in the in vivo forces. A more distinct trend appears when pedal forces are analyzed. In figure 5B the maximum pedal forces at different cadences were plotted for the average patient. It is shown, that the peak magnitude of F pi decreases as the cadence increases, ranging from 32% BW at 40 RPM to 23% BW at 90 RPM. Influence of seat height The ergometer cycling that was performed at specific powers by the patients at a normal saddle height was repeated at a lower saddle height (-9cm). In table 3, the values of the
in vivo hip joint force and the corresponding pedal forces for the average patient at different powers and different seat heights and the percentage increase are listed. The contact forces at low saddle height are slightly increased to the forces at normal saddle height, between 7-15%. The pedal forces at low saddle height are slightly higher (1-2% BW) than the pedal forces at normal saddle height.
Discussion The goal of this study was to analyze the acquired in vivo load data and external forces of a THA patient during cycling. The focus was on the effects of a sport especially because the profile of people receiving total hip replacements has changed over time. Influence of power Ergometer cycling at different power outputs has been extensively studied and various factors have been investigated; muscle activity, muscular power output (Ericson, 1988), pedal force application and in-shoe pressure distribution (Sanderson et al., 2000). In several studies, pedal forces have shown to increase with power, which was confirmed by the present study too. At higher power settings, positive hip extensor and flexor power output increases, the former at a higher rate than the latter (Ericson, 1988). Our results confirm that these effects can be translated to hip joint loads, since contact forces increased with higher power levels. Workload has been established as the most important variable to engender change in joint loads and muscular activity (Ericson et al., 1986). Influence of cadence
The power generated on an ergometer is the product of the crank torque and angular velocity. Therefore, to generate a specific power, the combinations of the applied force and the pedalling rate are important. Physical laws state that as cadence is increased, lesser force is required to maintain a constant power output. This effect holds for the pedal forces. At a constant power output of 90 W, pedal forces linearly decreased from 32 to 23% BW as cadence increased from 40 to 90 RPM. However this trend is not reflected in the in vivo loads at higher cadences. The contact forces at the joint do reduce as the cadence shifts from 40 to 60 RPM, however at 90 RPM a deviation is encountered, where the magnitude rises. Several studies have reported changes in muscle work (Neptune and Herzog, 1999), muscle co-ordination (Li and Caldwell, 1997; Neptune and Hull, 1996) and muscle activation dynamics (Neptune and Kautz, 2001) in the lower extremity during changes in cadences. Consequently, the changes in muscle activity have been well documented in cycling literature. However this study provides an insight into the effects of pedalling speed on in vivo joint forces. Previous studies on cadence effects can provide an explanation for the effect changing cadence is having on the hip joint loads. The increased magnitude of in vivo loads at higher pedalling speeds is the consequence of higher force generation by muscles and co-contraction of muscles. Activation dynamics come into play at higher cadences because of rapid contraction-relaxation cycles that muscles need to undergo. The delay between excitation and force generation, gives rise to force during the relaxation phase (lengthening of the muscle) rendering it useless; also called negative muscle work (Neptune and Herzog, 1999). To overcome this and maintain a constant power output, the muscles have to produce more positive work, leading to a higher contact force at the joint (Neptune and Herzog, 1999;
Neptune and Hull, 1996). Another cause is co-contraction of muscles at higher cadences, required for control and stabilization of joints. The preferred pedalling rate for cyclists is approximately 90 RPM and cadence effects usually begin after 90 RPM (Neptune and Hull, 1996). All previous studies have been done with healthy inexperienced cyclists. It should be noted that in this study, all participants are THA patients. The minimum in vivo contact forces were encountered at 60 RPM, therefore this cadence is recommended for rehabilitation. Due to the age, sub-optimal fitness levels and recovering status of patients, there were limitations on collecting data at higher cadences than 90 RPM. Influence of saddle height The effect of saddle height adjustments on pedal forces has previously been reported. Lower saddle heights produced higher resultant pedal forces but lower force effectiveness (Bini et al., 2011). In this study also, slightly higher pedal forces were encountered at low saddle height configuration at most cycling settings. Studies to date have included triathletes and cyclists with competitive experiences and the effects on joint angles and joint mechanical work have been collected or derived (Bini et al., 2011). Moreover, little research has been done regarding the hip joint, with more importance given to the knee joint. This study is the first to look into the saddle height effects on in vivo hip joint loads and uses data from cyclists with THA. The contact forces arising within the hip joint are slightly higher at a low saddle height; there is approximately an increase in contact forces between 7-15% on average. A study of muscle recruitment pattern during different saddle heights showed an increase in muscle activity with lower saddle heights. This trend was more prevalent in quadriceps and hamstrings (Raymond et al., 2005). The increased muscle forces at lower saddle
height are a possible explanation for increased hip joint contact forces. For hip arthroplasty
patients,
therapeutic
ergometer
cycling,
during
rehabilitation
is
recommended to be done at a normal saddle height as lower contact forces can be expected. Conclusion The in vivo measurements during cycling with instrumented hip implants have shown that comparing with other activities, cycling is an athletic activity with low-impact to the hip joint. During cycling inter-patient differences were observed in F res, this can be attributed to several factors like post-operative time, fitness level, agility and weight. For example, patient H9L was considerably heavier than the rest of the patients, which explain the high deviation of his contact forces from the rest of the patients. Additionally, there was a high inter-individual difference in post operative time; varying from 6 to 30 months. Besides the high inter-individual differences, there exists a wide range of contact forces even intra- individually, which can be a result of change in style and/or fatigue during the cycling measurement. When comparing the forces applied on both ipsilateral and contralateral pedals, it became undeniable that the operated and normal hip joints and surrounding muscles are not used equally in any patient. These inter-individual differences can be attributed to several factors like agility and athletic activities in daily living and also the patient’s adherence to the rehabilitation program. These factors ultimately affect the recovery of musculature around the hip joint. For example, a study compared the outcomes of patients receiving weight-bearing and postural stability exercises and patients doing basic isometric and active range of motion exercises 4 months after THA. The former
had significantly improved muscle strength, postural stability and self-perceived function (Trudelle-Jackson and Smith, 2004). Another factor to consider is whether the patient’s operated leg is the same as the dominant leg. All patients except H7R, had their dominant leg operated, i.e. their ipsilateral leg is their dominant leg. In figure 2B for patient H7R, the difference between the ipsilateral and contralateral pedal forces is greater than any other patient. Limitation The present study’s main objective is to measure the actual hip joint loads in vivo, which is accomplished with patients recovering from total hip replacements. The participants in this study are comparatively have lower fitness levels and are less experienced than the typical athletic cyclist is. Due to these reasons and the fact that the hip joints are different in the groups of participants (natural and artificial), limitations exist while correlating data. The small number of subjects participating in this study serves as a limitation.
Additional Data Selected examples of the in vivo measurements, on which this study was based, are published in the public data base www.orthoload.com.
Acknowledgements This work was supported by Deutsche Forschungsgemeinschaft (Be 804/19-1), the German Federal Ministry of Education and Research (BMBF 01EC1408A; OverloadPrevOP; SPO3), Deutsche Arthrose-Hilfe and the OrthoLoadClub.
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Figure Legends
Figure 1: In vivo hip joint loads and corresponding pedal forces; 90W 40 RPM; average subject
Figure 2: In vivo hip joint force (A) and pedal force during cycling; 40 RPM; 90W
Figure 3: Pattern of the in vivo hip joint force and pedal forces at 40 RPM and different ergometer power; average subject
Figure 4: Relation between maximum hip joint forces and the maximum ipsilateral pedal forces of each patient (coloured) and the average subject (black)
Figure 5: Maximum of the in vivo joint forces (A) and pedal forces (B); 90W;
Table 1:
Subjects investigated
Patient
H2R
H6R
H7R
H8L
H9L
Age
65
69
54
56
54
Body Weight
81
83
94
88
118
pOP Month
30
12
15
10
6
Site of THA
Right
Right
Right
Left
Left
Table 2:
Parameters adopted for cycling measurements
Saddle Height
Normal
Low [- 9cm]
Power at 40 RPM
50, 70, 90, 110 and 130W
50, 70, 90 and 110W
Cadence at 90W
40, 70 and 90 RPM
40, 70 and 90RPM
Table 3:
In vivo joint force and ipsilateral pedal forces at different seat heights and ergometer power; average subject
Ergometer
Normal Saddle Height
Low Saddle Height [- 9cm]
Δ Joint Force
Power
Joint Force
Pedal Force
Joint Force
Pedal Force
[W]
[% BW] (STD)
[% BW] (STD)
[% BW] (STD)
[% BW] (STD)
50
67 (13)
23 (3)
72 (13)
25 (3)
7
70
80 (14)
28 (3)
91 (16)
29 (3)
12
90
97 (16)
32 (3)
114 (21)
34 (3)
15
110
118 (19)
38 (3)
134 (20)
39 (3)
12
[%]