Hip abductor muscle activity in persons with a hip prosthesis while carrying loads in one hand.

Key Words: Electromyography, Hip abductors, Hip joint, Load carriage, Prosthesis.

Total hip arthroplasty is the most common major orthopedic procedure performed in the elderly population.[1] Due to the aging population, the number of persons requiring a hip prosthesis will increase in the next century. Although this operation is considered successful for most elderly people,[2,3] loosening of the femoral or acetabular component of the arthroplasty remains a major postoperative problem.[4]

Studies in the early 1980s showed that loosening of cemented femoral components occurred in 30% to 40% of patients 10 years after surgery.[5,6] Mechanical and biologic factors have been implicated as factors contributing to the high incidence of prosthetic loosening.[7-12] Mechanical factors include loads,[9,12] micromotion,[10,11]and characteristics of the implant materials. Biologic factors include bone remodeling,[8] sepsis,[8] and osteolytic responses to particulate debris from the implant.[7] Until sufficient long-term data emerge from clinical trials, the debate regarding the most effective and lasting method of hip implant fixation will continue.[13-20]

Physical therapists are actively involved in the physical rehabilitation of the person with a new hip prosthesis. Therapy includes gait training and exercise,[21-23] aerobic conditioning, and giving advice on means of hip joint "protection."[24] Protecting a prosthetic hip from unusually high and unnecessary forces may help retard premature loosening of the prosthetic hip, thereby increasing the implant's functional longevity. One common functional activity that likely places excessive force on the hip is carrying a load in a single hand.[25] The purpose of this study, therefore, was to record the surface electromyographic (EMG) activity of the hip abductor (HA) muscles to gain insights into the relative force demands that these muscles place on the prosthetic hips of persons who carry loads in one hand.

Background

To understand the proposed relationship between carrying a load and the assumed magnitude of forces produced across a prosthetic hip, the biomechanical role of the HA muscles can be considered.[25-28] The primary HA muscle is the gluteus medium muscle; however, the gluteus minimus muscle, the tensor fasciae latae muscle, and the anterior fibers of the gluteus maximus muscle also contribute to this action.[29] The main function of the HA muscles is to provide frontalplane stability for the hip during the single-limb support phase of walking. To achieve this stability, the HA muscles must produce a torque large enough to match the torque produced by body weight.

This balance of frontal-plane torques is shown in a static model that assumes the pelvis is held fixed and stationary over the femoral head component of the prosthetic hip during single-limb support (Fig. 1). Due to the difference in length of the moment arms used by the HA muscles and body weight (D versus[D.sub.1] in Fig. 1), the HA muscles must produce a force of about twice that of body weight to ensure frontal-plane equilibrium. The sum of this HA muscle derived force plus the force of body weight may reach 3 to 3.5 times body weight during mid-stance.[30] To achieve frontal-plane equilibrium while walking with a prosthetic hip, a "reaction force" must develop across the femoral and acetabular components of the prosthesis (Fig. 1). Based on this model, the force produced by the HA muscles is the largest contributor to the prosthetic hip reaction force. Reducing the need for excessive forces from the HA muscles should, in theory, minimize the forces produced across the hip and therefore reduce the potential for loosening of the prosthetic hip components.

Research on young subjects without hip disease has shown that surface EMG voltage from the HA muscles is dramatically influenced by the method of carrying a load in one hand.[31,32] The amount of normalized EMG activity produced by the HA muscles during the mid-stance phase of walking was assumed to reflect the relative muscular demands placed on the underlying hip joint. The data from these studies consistently showed that carrying a load in the hand opposite (de, contralateral to) to a given HA muscle produced greater EMG activity than walking without a load. Furthermore, carrying a load in one hand on the same side of the HA muscle (de, ipsilaterally) produced equivalent or less EMG activity than walking without a load. These results were attributed to the relative position and length of the external moment arm of the loads. The investigators in these studies concluded that as a method of hip joint protection, a person with hip disease should avoid or limit carrying loads in the hand contralateral to the affected hip. If a load must be carried in one hand, however, the load should be kept minimal and carried on the same side as the affected hip.

This study was performed to determine whether the above logic on hip joint protection can be applied to older individuals with hip prostheses. Electromyographic data from the HA muscles were analyzed as loads were carried in the hand ipsilateral and contralateral to the side of the prosthetic hip. This study focused on the biomechanical events during the mid-stance phase of walking. The normalized EMG data from the HA muscles were used as an index to reflect the presumed relative force generated by the HA muscles. Based on the model shown in Figure 1, changes in HA muscle EMG activity were assumed to be associated with similar relative changes in forces due to muscle activity across the prosthetic hip. Because the mathematical relationship between EMG activity and muscle force is not known for the HA muscle, no attempt was made to use EMG activity as an absolute measure of either muscle or joint force.

Method

Subject Selection Process

Twenty-five relatively active persons with a single hip prosthesis were selected for this study. Subjects were recruited through an advertisement placed in the local newspaper, regional hospitals, and arthritis support groups. At the time of the study, subjects lived in or near a moderately large metropolitan midwestern city in the United States. All subjects were paid for their time and received free consultation regarding their exercise program and suggestions on ways to minimize stress on their prosthetic hip.

Subjects selected for this study had to meet the following criteria. There could be only one prosthetic hip, and the implant could not be the result of a revised procedure. The operation must have been secondary to osteoarthritis. Subjects with a hip implant due to rheumatoid disease, avascular necrosis, or congenital dysplasia were not selected. The prosthetic hip also must have been the only joint prosthesis in the subject's body, and all other joints must have been free of joint disease. Several subjects selected for the study reported that they experienced joint pain that was minor but not sufficient to be considered disabling.

All subjects selected for this study stated that they were in good health and independent in all their activities of daily living. No subject reported respiratory disease, heart or vascular disease, or diabetes. Subjects did not require the use of an assistive device to aid in walking for distances less than 0.8 km (1/2 mile).

I performed a brief physical examination on and administered a questionnaire to each subject prior to his or her acceptance into the study. During the examination, major muscle groups were checked for weakness of the lower extremity (through manual resistance) and hip instability was tested. Gait abnormality, pain, or other conditions that might compromise subject safety during the experiment were noted. Subjects selected for the study did not require any special footwear or orthoses for walking.

Subject Profile

Nine women and 16 men were selected. All subjects signed consent forms. Subjects ranged in age from 40 to 86 years (X=63.7, SD=10.7), in weight from 498.2 to 1,085.4 N* (X=759.3, SD=163.7), and in height from 1.52 to 1.92 m (X=1.71, SD=0.10). Twelve subjects had the prosthetic hip on their right side, and 13 subjects had the prosthetic hip on their left side. The time since hip surgery ranged from 5 to 96 months (X= 24.9, SD = 21.4) .

Instrumentation

The EMG instrumentation used in this study has been described in earlier work.[31-33] Surface EMG data were collected from the HA muscles as subjects walked on an indoor, hard-surfaced walkway while carrying loads in one hand of varying weight up to 15% of their body weight. The EMG unit consisted of surface on-site electrodes, a ground electrode, an oscilloscope, a signal conditioning unit,[dagger] a personal computer, an analog-to digital converter, and software for data collection and data reduction. Raw bipolar EMG data were processed by using the root-mean-square (RMS) method by producing a linear envelope, or average voltage, over a specified time. The time constant used for the RMS processing was 55 milliseconds. The sampling rate of the processed EMG data was 100 times per second. The raw EMG signals were monitored for artifact by an oscilloscope.

Subjects wore footswitches attached to rubber galoshes placed over their shoes. The footswitches produced voltages that associated EMG voltage with a particular phase of gait. Two on-off footswitch closures defined the mid-stance phase of gait as the time interval between the instant of footflat and just prior to heel-off. Electromyographic and footswitch voltages traveled between subject and signal processor/computer via a 12.2-m (40-ft) cable.

Procedure

Preexperimental protocol. Subjects were led to a private room for the application of the EMG electrodes, ground plate, and rubber galoshes. The skin over the posterolateral gluteal region was cleansed with alcohol. An EMG electrode was placed on the skin superficial to the belly of the gluteus medius muscle as described in earlier work.[25,31] The ground electrode was placed over the anteromedial aspect of the tibia on the side of the prosthetic hip. Electrode placement was verified by palpation of the gluteus medius muscle during isometric contraction and observation of the raw EMG signal as the subjects stood in single-limb support on the side of the active muscle.

Subjects were next taught to walk at a relatively constant self-selected walking speed. Pilot work indicated that the natural free walking speed of many subjects exceeded the speed that the heavier loads could be comfortably carried. To reduce the self-selected walking speed, subjects were instructed to walk using a cane held in the hand opposite their prosthetic hip. Using the cane had the effect of reducing walking speed to a level that coincided with a tolerable speed for carrying the heaviest loads. After the walking speed was established, the cane was no longer used in this experiment.

After at least 3 minutes of walking, the subjects' average walking speed was determined by stopwatch to the nearest 10th of a second (X=0.82 m/s, SD=0.09 m/s, range=0.61-1.00 ms, for all 25 subjects). Each subject's average walking speed was determined over three trials of walking a 10-m distance. Subjects repeated all subsequent walking trials at a speed within 10% of their own self-selected target speed. All subjects were able to learn to control their walking speed within the required range.

Subjects next practiced walking in a natural manner with the instrumentation in place. As subjects performed these preliminary walks, all data channels were sampled and reviewed on the computer screen. After confirmation of proper function of the instrumentation, the gain to the EMG amplifiers was set so that the EMG voltage produced during a maximal isometric contraction of the HA muscles was well under the maximal limits expected by the computer's analog-to-digital converter.

Subjects were asked to carry a load the way a person carries a suitcase. The loads consisted of weights placed in a plastic container with a hinged handle, with dimensions of 18 x 20 x 23 cm (7 x 8 x 9 in). The total weight of each load was adjusted to 5%, 10%, and 15% of each subject's body weight.

Experimental protocol. Before the actual start of the load-carrying experiments, a preexperimental EMG baseline was established for each subject. This baseline was determined by averaging the HA muscle EMG voltage produced during the mid-stance phase of walking without a load. For each no-load trial, the sampling of EMG data began as the subject walked across a 2-m mark on the walkway and continued for 10 seconds. Subjects were instructed to stop walking after the 10second period. An EMG no-load baseline voltage was determined by averaging these data across the four walking trials.

During the load-carrying experiments, data were collected while each subject carried three different loads (de, 5%, 10%, and 15% of body weight) on each side of the body (de, ipsilateral and contralateral to the side of the prosthetic hip). The order of each of the six loadcarrying conditions was randomized. Subjects were allowed a 90-second rest between walking trials as the experimenters displayed and verified the footswitch and EMG pattern on the computer screen. Data were accepted for analysis only after the target walking speed was confirmed and a typical footswitch pattern was displayed on the computer screen. Before data collection, subjects performed a practice walk for each side of carrying.

Data were collected during the load-carrying trials in a manner similar to that described for the preexperimental no-load tests. One difference, however, was that data for only two walking trials were collected for each of the six load-carrying conditions. This experimental design provided data on approximately eight complete walking cycles per walking trial. On average, each of the six load-carrying conditions produced data that were averaged over 16 complete gait cycles per subject. Electromyographic voltages produced as subjects carried a load were normalized to a percentage of the preexperimental voltage baseline. This normalized EMG value was expressed as a percentage of the baseline activity (% EMG).

Reliability Assessment

Following the load-carrying phase of the experiment, each subject reestablished a postexperimental no-load EMG baseline by repeating the preexperimental no-load walking trials. To determine the intrasubject reliability of the EMG measurements, a comparison was made between the grand mean EMG voltage (in millivolts) produced in the preexperimental no-load walking trials and the EMG voltage produced during the postexperimental no-load trials. Approximately 3 hours of time separated these two measurements. Each grand mean was calculated by averaging all 25 subjects' EMG voltage from the side of the prosthetic hip during the mid-stance phase. The mean preexperimental no-load EMG voltage was 149.2 mV, and the mean postexperimental no-load EMG voltage was 143.1 mV. This 4% difference in EMG baseline was considered insignificant. This difference would not have a systematic affect on the results of this study because the order of the performance of the load-carrying conditions was completely random. An intraclass correlation coefficient (ICC[1,k]) of .991 was calculated for the preexperimental and postexperimental data (P<.0001).[34,35]

Data Analysis

The complete data for all 25 subjects consisted of normalized %EMG values obtained from the HA muscles on the side of the prosthetic hip for the six loadcarrying conditions. All %EMG data were collected during the mid-stance phase of walking. The mean %EMG value for each of the six load-carrying conditions was based on a grand mean of approximately 16 complete gait cycles per subject, averaged over all 25 subjects.

An analysis of variance (ANOVA) with a repeated-measures design was performed on the normalized %EMG data.[34] The dependent variable was HA muscle %EMG produced on the side of prosthetic hip. The independent variable was load-carrying condition (de, three load weights x two sides of carrying). The %EMG measures for the three weights for each side of carrying were compared against each other and then against 0% (de, the preexperimental no-load EMG value) by using a multiple t tests with Bonferroni adjustments.[31,34] These adjustments maintained the a priori alpha level by dividing .05 by the number of preplanned comparisons. Comparisons between %EMG means were considered as different only when the associated probability value for each comparison was less than the adjusted alpha level.

Results

An ANOVA test on the mean %EMG produced by the HA muscles showed a main effect for the variable "load-carrying condition" (F=108.7, P<.0001). The Table shows the descriptive data for the six load-carrying conditions averaged over all 25 subjects. The %EMG means for the six load-carrying conditions are depicted in Figure 2.

Table
Descriptive Statistics for Hip %EMG(a) Data for Six Load-Carrying
Conditions (N=25)

Load Condition(b)       X          SD       Minimum      Maximum

Ipsilateral
  5% of BW(c)        -10.64       12.65      -48.0         11.0
  10% of BW          -16.92       19.46      -83.0          7.0
  15% of BW          -16.96       15.50      -53.0          9.0

Contralateral
  5% of BW            16.76       15.53      -6.0          65.0
  10% of BW           38.88       20.42       5.0          73.0
  15% of BW           58.44       26.60       9.0         113.0

(a) % EMG=percentage of electromyographic voltage during no-loaf
walking. Negative values indicate EMG less than that produced
walking without a load.
(b) "Ipsilateral" and "contralateral" refer to the side on which
the load was carried relative to the hip abductor muscles and the
prosthetic hip.
(c) BW=body weight.

Each of the six load-carrying conditions produced an HA muscle %EMG that was different from 0% (de, the EMG activity produced while walking without a load). Furthermore, the mean %EMG values for the three ipsilateral load-carrying conditions were equivalent to each other. In contrast, the mean %EMG values for the three contralateral load-carrying conditions were different from each other.

Discussion

Carrying weights of up to 15% of body weight in a position contralateral to the prosthetic hip produced an HA muscle %EMG greater than that produced while walking without a load (Fig. 2). The larger the contralaterally held load, the greater the demands placed on the HA muscles and, in theory, on the prosthetic hip.

The reason for such a large %EMG demand on the HA muscles while carrying the contralaterally held load becomes clear in the model shown in Figure 3A. Assuming static equilibrium about the right prosthetic hip, the right HA muscles must generate a relatively large counterclockwise torque equal to the comoined clockwise torques produced by the load and body weight. Large muscular forces are required because the moment arm of the HA muscle is much smaller than the moment arm caused by carrying a load on the contralateral side (compare D and [D.sub.2] in Fig. 3A). To ensure frontal-plane equilibrium during mid-stance, the HA muscles must generate a very large force that crosses between the prosthetic components.

The magnitude of the reaction force at the prosthetic hip may be very large while carrying a contralaterally held load weighing 15% of body weight. Figure 3B provides the data needed to estimate both the HA muscle force and the prosthetic hip reaction force.[36,37]

The frontal-plane model in Figure 3 assumes that the right hip is in static equilibrium. For simplicity, the model does not include dynamic variables associated with the mid-stance phase of walking, or forces and torques produced outside of the frontal plane. The model, therefore, would contain error in predicting absolute force and torque magnitudes. I believe that the static model is nevertheless a useful tool for explaining the nature of the %EMG responses of the HA muscle for carrying loads in one hand. As shown in Figure 3, dividing the sum of the clockwise torques (dashed circles) by the HA muscle moment arm (de, 0.0439 m) yields 2,193.6 N (493.2 lb) of HA muscle-generated forces. For the hip to remain stable during mid-stance, the prosthetic hip reaction force must be equal to the sum of the HA-derived force, the force of body weight, and the weight of the contralaterally held load. Using the data supplied in Figure 3B, the approximate prosthetic hip reaction force is estimated to be 2,952.9 N (663.9 lb), or about four times body weight.

Physical therapists and others need to be aware of the presence of such high forces,[21] especially during the early postoperative phase with cementless prosthetic hips. Increased reaction forces across the prosthetic hip may contribute to increased wear and the subsequent production of particulate debris from the implant. This debris may cause increased macrophage activity within bone, which may add to the osteolysis and subsequent loosening of the prosthesis.[7]

Carrying the three weights on the side ipsilateral to the prosthetic hip produced dramatically less HA muscle %EMG than that produced with the contralaterally held loads (Fig. 2). The reason for the decrease in HA muscle %EMG (and assumed prosthetic hip reaction force) while carrying ipsilaterally held loads can be at least partially explained by Figure 4A. Carrying the load ipsilateral to the prosthetic hip markedly reduces the torque needed by the right HA muscles for frontal-plane equilibrium. The ipsilaterally held load produces a counterclockwise torque that is in the same rotary direction as that required by the HA muscle to balance the torque due to body weight. In essence, the weight of the ipsilaterally held load shares the torque responsibility of the HA muscle, thereby minimizing the force component created by the HA of the total prosthetic hip reaction force. The calculations shown in Figure 4B indicate that carrying a load weighing 15% of body weight ipsilateral to a prosthetic hip requires only 808.7 N (181.8 lb) of HA muscle force and 1,568 N (352.5 lb) of prosthetic hip reaction force. This reaction force is only about half the force estimated for carrying a contralaterally held load weighing 15% of body weight.

The HA muscle %EMG values produced while carrying the ipsilaterally held loads were less than the EMG values generated while carrying no hand-held load (Fig. 2). Calculations using the static equilibrium model suggest that the reaction force from an ipsilaterally held load would also be less than the no-load condition. By using the data and equations supplied in Fig. 4B (but ignoring the ipsilaterally held load and [D.sub.3] variables), calculations show the prosthetic hip reaction force for the no-load carry was 1,916.5 N (430.9 lb). The reaction force for the ipsilaterally held load weighing 15% of body weight was estimated to be only 1,568 N (352.5 lb), which was about 20% less than that produced during the no-load condition. This reduction is due to the decreased torque demand on the HA muscle during mid-stance.

The data from this study indicate that carrying loads in the hand ipsilateral to the side of a prosthetic hip will lead to less force on the underlying prosthetic components than carrying loads in the hand opposite the prosthesis. This finding does not imply that persons with a prosthetic hip should be encouraged to carry loads in this fashion. In contrast, as a general rule, I believe that persons with a prosthetic hip should be discouraged from carrying loads on one side of the body, except when absolutely necessary. An ipsilaterally held load protects the prosthetic hip only at the expense of higher HA forces on the opposite hip.[32] For example, a load held ipsilaterally in the right hand protects the right prosthetic hip while in right mid-stance, but this same load becomes a contralaterally held load for the left hip during mid-stance. If the left hip is healthy, this increased demand may be an acceptable trade-off. Persons with bilateral prosthetic hips, or a prosthesis and some pathology of the opposite hip, must consider dividing a load into two loads and using both hands, or use a mechanical means of transport.[31]

Researchers[25,31,32] have measured the HA muscle %EMG responses to carrying loads using relatively young persons with asymptomatic hip joints. In one study,[32] the HA muscle %EMG was recorded as subjects with a mean age of 21.5 years carried loads of between 3% and 30% of body weight in one hand. In general, when comparing identical load weights, the %EMG responses obtained in the present study were very similar to the EMG responses obtained from the young subjects with normal hip joints (Fig. 5). This similarity in EMG responses presumably would continue for loads that weigh greater than 15% of body weight. If so, subjects with a prosthetic hip joint would experience very large %EMG responses and assume very large and potentially damaging reaction forces.

Summary and Conclusions

The EMG data from this study and from other publications[25,31,32] suggest that the method of carrying loads in the hand affects the force demands on the HA muscles and the underlying joints. The results show an increased %EMG response from subjects' HA muscles during the mid-stance phase of walking while carrying loads in the hand contralateral to their prosthetic hip. The assumption was made that the increased HA muscle %EMG was associated with increased reaction forces at the prosthetic hip.

Subjects showed a marked reduction in HA muscle %EMG during mid-stance while carrying loads in the hand ipsilateral to the side of their prosthetic hip. On average, an ipsilaterally held load of 15% of body weight produced about 17% less EMG activity than that produced while walking without a load. The reduced %EMG was assumed to be associated with decreased reaction forces at the prosthetic hip.

Whether the differences in suspected hip forces from the two different ways of carrying loads actually lead to loosening of the implant cannot be determined from this study. The data, however, suggest that therapists and physicians should advise patients of the possible deleterious consequences of subjecting their hip prosthesis to the larger forces that are likely to occur while carrying a load on the side contralateral to their implant.

Acknowledgments

I am grateful to the following persons for their assistance with subject recruitment, data collection, graphics, or statistical support: Tony Hornung, PT, Thomas Cook, PhD, PT, Richard Shields, PhD, PT, John Rosecrance, PhD, PT, Richard Jensen, PhD, PT,Joan Holcomb, and Gregg Furhman, PT.

[Figures 1 to 5 ILLUSTRATION OMITTED]

(*) 4.448 N= 1 lb. ([dagger]) Therapeutics Unlimited, 2835 Friendship St, Iowa City, IA 52240.

References

[1] Sharkness CM, Hamburger S, Moore RM, Kaczmarek RG. Prevalence of artificial hip implants and use of health services by recipients. Public Health Rep. 1993;108:70-75. [2] Pettine KA, Aamlid BC, Cabanela ME. Elective total hip arthroplasty in patients older than 80 years of age. Clin Orthop. 1991;266:127-132. [3] Gogia PP, Christensen CM, Schmidt C. Total hip replacement in patients with osteoarthritis of the hip: improvement in pain and functional status. Orthopedics. 1994;17:145-150. [4] Harris WH. The first 32 years of total hip arthroplasty. Clin Orthop. 1992;274:6-10. [5] Stauffer RN. Ten-year follow-up study of total hip replacement. J Bone Joint Surg [Am]. 1982;64:983-990. [6] Sutherland CJ, Wilde AH, Borden LS, Marks KE. A ten-year follow-up of one hundred consecutive Muller curved-stem total hip-replacement arthroplasties. J Bone Joint Surg [Am]. 1982;64:970-982. [7] Schmalzried TP, Kwong LM, Jasty M, et al. The mechanism of loosening of cemented acetabular components in total hip arthroplasty. Clin Orthop. 1992;274:60-78. [8] Volz RG, Wilson RJ. Factors affecting the mechanical stability of the cemented acetabular component in total hip replacement. J Bone Joint Surg [Am]. 1977;59:501-504. [9] Jones PR, Hukins DWL, Porter ML, etal. Aseptic loosening of the femoral component in cemented hip replacement. J Biomed Eng. 1992;14:379-384. [10] Markolf KL, Amstutz HC. Compressive deformations of the acetabulum during in vitro loading. Clin Orthop. 1983;173:284-292. [11] Markolf KL, Amstutz HC, Hirschowitz DL. The effect of calcar contact on femoral component micromovement. J Bone Joint Surg [Am]. 1980;62:1315-1323. [12] Wroblewski BM, Lynch M, Atkinson JR, et al. External wear of the polyethylene socket in cemented total hip arthroplasty. J Bone Joint Surg [Br]. 1987;69:61-63. [13] Barrack RL, Mulroy RD, Harris WH. Improved cementing techniques and femoral component loosening in young patients with hip arthroplasties: a 12-year radiographic review. J Bone Joint Surg [Br]. 1992;74:385-389. [14] Harris WH. Total hip replacement: "cement versus cementless" resolution. Bull Rheum Dis. 1994;43:1-4. [15] Morrey BF, Kavanagh BF. Cementless joint replacement: current status and future. Bull Rheum Dis. 1987;37:1-7. [16] Davey JR, Harris WH. A preliminary report of the use of a cementless acetabular component with a cemented femoral component. Clin Orthop. 1989;245:150-155. [17] Schmalzried TP, Harris WH. Hybrid total hip replacement: a 6.5-year follow-up study. J Bone Joint Surg [Br]. 1993;75:608-615. [18] Wroblewski BM. Cementless versus cemented total hip arthroplasty: a scientific controversy? Orthop Clin North Am. 1993;24:591-597. [19] Wixson RL, Stulberg SD, Mehlhoff M. Total hip replacement with cemented, uncemented, and hybrid prostheses. J Bone Joint Surg [Am]. 1991;73:257-270. [20] Evans BG, Salvati EA, Huo MH, Huk OL. The rationale for cemented total hip arthroplasty. Clin Orthop. 1993;24:599-610. [21] Krebs DE, Elbaum L, Riley PO, et al. Exercise and gait effects on in vivo hip contact pressures. Phys Ther. 1991;71:301-309. [22] Strickland EM, Fares M, Krebs DE, et al. In vivo acetabular contact pressures during rehabilitation, part 1: acute phase. Phys Ther. 1992;72:691-699. [23] Givens-Heiss DL, Krebs DE, Riley PO, et al. In vivo acetabular contact pressures during rehabilitation, part II: postacute phase. Phys Ther. 1992;72:700-705. [24] Neumann DA. Biomechanical analysis of selected principles of hip joint protection. Arthritis Care and Research. 1989;2:146-155. [25] Neumann DA, Cook TM. Effect of load and carry position on the electromyographic activity of the gluteus medius muscle during walking. Phys Ther. 1985;65:305-311. [26] Inman VT. Functional aspects of thc abductor of thescles of the hip. J BoreJoint Surg. 1947;29:607-619. [27] McLeish RD, Charnley J. Abduction forces in the one-legged stance. J Biomech. 1970;3:191-209. [28] Maquet P. Biomechanics of the Hip: As Applied to Osteoarthritis and Related Conditions. New York, NY: Springer-Verlag New York Inc; 1985. [29] Dostal WF, Soderberg GL, Andrews JG. Actions of hip muscles. Phys Ther. 1986;66:351-359. [30] Rydell N. Forces acting on the femoral head-prosthesis: a study on strain gauge supplied prosthesis in living persons. Acta Orthop Scand. 1966;37(suppl 88):1-132. [31] Neumann DA, Cook TM, Sholty RL, Sobush DC. An electromyographic analysis of hip abductor muscle activity when subjects are carrying loads in one or both hands. Phys Ther. 1992;72:207-217. [32] Neumann DA, Hase AD. An electromyographic analysis of hip abductors during load carriage: implications for hip joint protection. J Orthop Sports Phys Ther. 1994;19:296-304. [33] Neumann DA, Soderberg GL, Cook TM. Electromyographic analysis of hip abductor musculature in healthy right-handed persons. Phys Ther. 1989;69;431-440. [34] SAS for the Personal Computer. 7th ed. Cary, NC: SAS Institute Inc; 1987. [35] Shrout PE, Fleiss J. Intraclass correlations: uses in assessing rater reliability. Psychol Bull. 1979;86:420-428. [36] Neumann DA, Soderberg GL, Cook TM. Comparison of maximal isometric hip abductor muscle torques between hip sides. Phys Ther. 1988;68:496-502. [37] Olson MA, Smidt GL,Johnston RC. The maximal torque generated by the eccentric, isometric, and concentric contractions of the abductor muscles. Phys Ther. 1972;52:149-157.

DA Neurmann, PhD, PT, is Associate Professor, Department of Physical Therapy, Marquette University, Box 1881, Milwaukee, WI, 53201-1881 (USA) ([e-mail]).

This study was approved by the Human Subjects Committee at Marquette University.

This project was funded by a grant from the National Arthritis Folmdation and by Marquette University and Columbia Hospital's Physical Therapy Department and Mttsculoskeletal Institute, Milwaukee, Wis.

Parts of this article were presented at the Joint (.ongress of the American Physical Therapy Nssociation-Canadian Physiotherapy Association; June 4-8, 1994; Toronto, Ontario, Canada.