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The Effect of Four Prosthetic Feet on Reducing Plantar Pressures in Diabetic Amputees

Sharon Hayden, PT
Reese Evans, CPO
Thomas G. McPoil, PhD, PT, ATC
Mark W. Cornwall, PhD, PT, Cped
Laurie Pipinich, CP

ABSTRACT

The purpose of this study was to determine the effect of four different prosthetic foot designs on the pattern of plantar pressure in the sound limb of unilateral transtibial diabetic amputees during walking. Twelve male community ambulators with a mean age of 56.3 years and a history of diabetes served as subjects. Pressure data were collected using the EMED Pedar in-shoe measurement system as subjects walked with the following four prosthetic feet: 1) Seattle Lite, 2) molded solid ankle, cushioned heel (SACH), 3) Sure-Flex III, and 4) multi-axial Greissinger. The measured results indicated that, except for the heel region, the type of prosthetic foot used did not significantly affect the peak pressures acting on the plantar surface of the intact limb. None of the prosthetic feet tested had a significant effect on reducing the pressure-time integrals for any of the seven plantar regions studied.

Key Words: diabetic, transtibial amputee, plantar pressure, prosthetic feet

Introduction

Lower extremity amputation is a frequent complication of diabetes that can result from vascular and neuropathic complications associated with the disease. It has been estimated that 40,000 amputations are performed annually in the United States on individuals with diabetes.¹ Although the rate of amputation in diabetic patients is approximately 15 times higher than in the nondiabetic population, the incidence of the remaining leg being amputated within the first 4 years after the loss of the first leg exceeds 50%.^2,3 Often, an ulceration on the plantar surface of the foot is the first insult that can initiate the series of events leading to eventual lower extremity amputation. Plantar ulceration has been linked previously to high plantar pressures in both retrospective and prospective investigations.^4,5,6,7 Thus, a strong association has been established between the plantar regions of high pressure and ulceration on the bottom of the foot.

A major concern for the unilateral diabetic amputee is the possibility of increased levels of plantar pressures acting on the remaining limb after amputation. Unfortunately, the extensive degree of gait training that is required to effectively and efficiently use the prosthesis after amputation can lead to increased loading on the foot of the remaining limb and possibly increased plantar pressures. Considering the high rate of remaining limb amputations in unilateral diabetics, the importance of selecting the most appropriate prosthetic components becomes critical. Pinzur et al. evaluated the magnitude of plantar pressures acting on the foot of the remaining limb of seven transtibial amputees who were diagnosed with peripheral vascular disease. Using seven discrete sensors positioned over the hallux, forefoot, and heel, their results suggested that prosthetic usage did not result in increased loading of the residual limb.⁸ In a more recent study, Veves et al measured the pressures under the remaining foot of diabetic amputees and compared them with the foot pressures of age- and weight-matched nondiabetic amputees, nonamputee diabetic patients with similar neuropathic involvement, and nondiabetic control subjects. Each group consisted of 11 subjects, and pressures were measured using an optical pedobarograph. The authors reported that the highest peak plantar pressures were present under the remaining foot in diabetic amputees and that these high pressures were associated with diabetic neuropathy.⁹

With recent advances (such as energy-storing properties) in the design of prosthetic feet, manufacturers have suggested that the use of these feet could possibly decrease the stresses being applied to the remaining limb of the diabetic amputee. Powers et al. demonstrated that prosthetic foot design had a significant influence on the loading of the sound limb in traumatic transtibial amputees.¹⁰ Using a force platform to determine if the same effects were observed in dysvascular transtibial amputees, Snyder et al. studied the effect of five different prosthetic feet on the loading of the sound limb in seven diabetic transtibial amputees. Their results indicated that the sound limb in the group of amputees had increased vertical force loading. In addition, certain prosthetic foot designs were shown to decrease the magnitude of the vertical forces acting on the sound limb.¹¹

Although the design of the prosthetic foot would appear to influence the vertical forces acting on the sound limb of the diabetic amputee, to date no study has investigated the effect of prosthetic foot design on the pattern of plantar pressures in this patient population. The purpose of this study was to determine the effect of four different prosthetic foot designs on the pattern of plantar pressures acting on the remaining limb of diabetic, unilateral transtibial amputees during walking. To assess both the magnitude and duration of the plantar pressures acting on the remaining limb in these amputees, both peak pressure and pressure-time integral were evaluated for seven regions on the bottom of the foot.

Methods

Subjects

Twelve male, diabetic patients with unilateral transtibial amputations were recruited to serve as subjects for the study. The subjects ranged in age from 42 to 72 years, with a mean age of 56.25 years. All subjects were recruited from the Maricopa Medical Center Amputee Program, and each met the following inclusion criteria: 1) a diagnosis of diabetes based on a medical history of elevated serum glucose levels; 2) an intact remaining limb with no current foot ulcerations; 3) a community ambulator requiring only the use of a cane during walking; 4) volume stability of the residual limb for at least 4 months before the start of the study; and 5) no current complications with the residual limb. Before the study began, all participants were screened to determine if protective sensation was present in the intact limb. Subjects with the inability to sense a Semmes-Weinstein 5.07 monofilament on the plantar surface of the sound, intact limb were classified as insensate. Based on this testing, 8 subjects were classified as insensate and 4 subjects were classified as sensate. Table 1 presents the demographics of these study participants. The Maricopa Medical Center Human Subjects Review Committee approved the study protocol. All subjects provided informed, written consent before participating in the investigation.

Foot Selection and Prosthetic Design

The following prosthetic feet were tested in random order: 1) the Seattle Lite foot (Model #SLF 135, 3/8-inch heel; Model Instrument Development, Poulsbo, WA); 2) the molded solid ankle, cushioned heel (SACH) foot (Model #SFP, 3/8-inch heel, built-in pyramid; Ohio Willow Wood, Mount Sterling, OH 43143); 3) the Sure-Flex III foot (Model #0145-174, built-in titanium pyramid; Flex-Foot, Inc., Aliso Viejo, CA); and 4) the multi-axial Greissinger foot (Model #1AI3, 25-mm heel, enclosed ankle joint; Otto Bock Orthopedic Industry, Inc., Minneapolis, MN). Based on the Durable Medical Equipment Regional Carriers (DMERC) Region D Supplier Guidelines at the time of data collection, these four feet were classified as either level 1 and 2 (nonenergy-storing) feet or level 3 (energy-storing) feet.¹² (DMERC is a group of four insurance companies selected by the Health Care Financing Administration to process durable medical equipment, prosthetics, orthotics, and supplies claims for the Medicare program.) The molded solid ankle, cushioned heel foot (SACH) was categorized as level 1, the multi-axial Greissinger foot (MULTI) as level 2, and the Seattle Lite (SEATTLE) and Sure-Flex III (FLEX) feet as level 3. All prostheses used in the current study were fabricated using an endoskeletal design with a total-contact, patellar tendon-bearing socket. The suspension systems used by the amputees were varied and included sleeve, pin, or supracondylar-suprapatellar designs.

Instrumentation

The EMED Pedar pressure measurement system (Novel Electronics, Inc., Minneapolis, MN), which has a sampling rate of 50 Hz, was used to record pressure and force data. The validity of the capacitance sensor used with the Pedar system has been documented previously.¹³ The EMED capacitance sensor insole was attached to the Pedar system. Consisting of a matrix of 90 to 100 capacitance transducers, the insole was approximately 2 mm thick. Before data collection began, both pairs of insoles were calibrated using a rubber bladder that was pressurized with compressed air. The calibration procedure involved applying a linear range of pressures to each insole, with a minimum pressure of 0 kPa and a maximum pressure of 700 kPa. The input pressure saturation range for the capacitance sensors used in the Pedar insole was 1200 kPa. In the current study, none of the subject trials achieved pressure saturation levels. The Pedar unit was connected to a computer by a 6-m cable for data collection and storage.

Procedures

After the informed consent was signed, the height and weight of each subject were measured and recorded. The fit and alignment of each tested prosthetic foot were reviewed and clinically optimized by a team of two certified prosthetists. Once proper alignment was verified, the subject donned first a pair of standardized, white cotton socks issued by the investigators, then his own footwear. To become familiar with the prosthetic foot and to ensure proper prosthetic alignment, each subject was instructed to walk 90 m, with rest periods as necessary. Once the amputee was familiarized with the selected prosthetic foot and the prosthetic alignment was judged by the certified prosthetists to be optimal, the EMED capacitance sensor insole was placed in the shoe of the sound limb. Pressure data were collected as each subject twice walked 16 m at his own, self-selected speed. Once data collection was completed, the subject was seated and his shoes and the prosthesis were removed. The subject was then allowed to rest for 15 minutes during the attachment of the next prosthetic foot to be tested. The same testing procedure described above for the first prosthetic foot was repeated for each of the remaining three prosthetic feet designated for testing. The order of testing for the four prosthetic feet was randomly determined. After completion of the study, each subject stated their preference regarding which prosthetic foot he would choose for permanent use.

Data Analysis

The EMED Pedar expert software system (Novel Electronics, Inc.) was used to select four consecutive steps from each of the two data collection trials for a total of eight steps. The steps were selected from the middle, 10-m portion of the 16-m walking distance to ensure that the subject was neither accelerating nor decelerating from his self-selected walking speed. This particular number of steps was selected for analysis based on the previous work of Kernozek et al.,¹⁴ who reported that a maximum of eight steps was needed to achieve an excellent level of reliability (> 0.90) using the EMED Pedar system for the variables of peak pressure and pressure-time integral.

Once the eight steps were selected, the EMED NovelWin software program (Novel Electronics, Inc.) was used to divide each step into seven regions, including the heel, midfoot, medial forefoot, central forefoot, lateral forefoot, hallux, and toes. For all subjects, these regions were consistently defined as a percentage of the total foot length and width of the subject's footprint, as illustrated in Figure 1 . The heel comprised the first 0% to 30% of foot length, the midfoot the next 30% to 60%, the forefoot the following 60% to 85%, and the hallux/toes the remaining 85% to 100%. The forefoot region width was divided into equal thirds, creating three forefoot regions. The hallux/toe region width was also divided into two parts, with the hallux region occupying the medial 40% and the toe region occupying the lateral 60%. Once the seven regions for each step were defined, peak pressure (in N/cm²) and pressure-time integral (in N/cm²-seconds) were calculated for each region. The variables peak pressure and pressure-time integral were not normalized in relationship to the subject's body, because previous research involving a diabetic population has shown that elevated body weight appears not to result in increased plantar pressures.¹⁵

Statistical Analysis

In addition to descriptive statistics, type (2,1) intraclass correlation coefficients (ICC) were calculated to determine the consistency of stance-phase duration for the eight steps collected for the foot of each subject's intact limb. The level of reliability for the intraclass correlation coefficients was classified according to the characterizations reported by Landis and Koch.¹⁶ These characterizations were designated as "slight" (if the coefficient ranged from .00 to .20), "fair" (if the coefficient ranged from .21 to 40), "moderate" (if the coefficient ranged from .41 to .60), "substantial" (if the coefficient ranged from .61 to .80), and "almost perfect" (if the coefficient ranged from .81 to 1.00).

A repeated-measures analysis of variance (ANOVA) was used to evaluate differences among the four prosthetic feet for the dependent variables peak pressure and pressure-time integral. A Tukey's post hoc comparison was used to determine differences among the conditions tested, with an alpha level of .05 used for all tests of statistical significance.

Results

The means and standard deviations for stance-phase duration were (in milliseconds) 928 ± 28 for SACH, 924 ± 25 for FLEX, 940 ± 25 for MULTI, and 951 ± 28 for SEATTLE. The results of the ICC for stance-phase duration were .903 for SACH, .890 for FLEX, .861 for MULTI, and .819 for SEATTLE. Listed in Table 2 are the means and standard deviations for the peak plantar pressure. The results of the repeated-measures ANOVA on the dependent variable peak pressure were not significant for any of the four prosthetic feet tested across all seven plantar regions, with the exception of the heel region. Based on results of the post hoc comparisons, the plantar pressures acting on the heel region for the FLEX prosthetic foot were significantly (p < .05) less than with the other three prosthetic feet studied.

The means and standard deviations for the pressure-time integral are listed in Table 3 . The results of the repeated-measures ANOVA on the dependent-variable pressure-time integral were not significant for any of the four prosthetic feet tested across all seven plantar regions.

Discussion

The first step in interpreting the results of this study was to determine whether the duration of stance phase was consistent among the eight steps collected from the foot of the intact limb of each subject. The consistency of the stance-phase duration is an important issue because previous studies have noted that changes in the walking cadence between trials can influence the duration of stance phase, which in turn can affect force and pressure values.^14,17 Because the values for the ICC were all above .81, the characterization of the stance-phase duration ICC would be classified as "almost perfect" for all four prosthetic foot groupings, according to the criteria established by Landis and Koch.16 Based on the ICC obtained, the authors concluded that the stance-phase durations were very consistent and that further analyses of the force and pressure data could be continued.

The purpose of this study was to determine the effect of four different prosthetic foot designs on the pattern of plantar pressures in the sound limb of diabetic, unilateral transtibial amputees during walking. Previous investigators have demonstrated that the design of the prosthetic foot could possibly influence the vertical forces acting on the sound limb of the diabetic amputee, but the effect of prosthetic foot design on the pattern of plantar pressures has not been shown previously.¹⁰ To assess the magnitude of pressures acting on the plantar surface of the foot of the intact limb, peak pressures were determined for seven plantar regions. In this group of amputees, the location of highest peak pressure within the shoe was the heel. In the forefoot region, the medial forefoot had the highest recorded peak pressures, followed by the central forefoot, the lateral forefoot, and the hallux. Although Pinzur et al. reported peak pressure values for the remaining limb of the seven amputees in their study, they only presented data obtained when the amputees were not wearing a shoe.⁸ The location pattern of the highest peak pressures in their diabetic amputees, however, followed the same pattern that was observed in the current study, even though all pressure data were collected when the amputees were wearing shoes. The results of the statistical analysis on peak pressure indicated that the only region with a significant reduction in pressure as a result of using different prosthetic feet was the heel. The reduction in the peak pressures acting on the plantar surface of the heel region occurred only when using the FLEX prosthetic foot. None of the four prosthetic feet, however, caused a significant change in the magnitude of peak pressures acting on any of the forefoot regions or the hallux. A reduction of peak pressure in these regions would be of greater interest because previous studies have documented that the most common sites of plantar ulceration in the diabetic patient with neuropathy are the forefoot and hallux.^5,18,19

To study the duration of the pressures acting on the plantar surface of the intact limb in the group of amputees studied, the pressure-time integral was also evaluated for seven regions on the bottom of the foot. Similar to the findings for peak pressure, the greatest pressure-time integral values were located in the heel, followed by the medial forefoot, the central forefoot, and the lateral forefoot. The results of statistical analysis indicated that none of the four prosthetic feet evaluated in the current study caused a significant reduction in the pressure-time integrals for any of the seven plantar regions. These findings suggest that none of the four prosthetic feet studied in this investigation significantly altered the duration of pressure acting on any of the seven plantar regions.

There are numerous factors that the prosthetist must consider when selecting the appropriate prosthetic foot for any unilateral transtibial amputee. These factors include the individual's potential level of activity, body mass, physical endurance, and desire to learn how to effectively use the device. Another concern that must be considered for the diabetic, neuropathic amputee is whether increased stresses will be placed on the intact side as a result of prosthetic foot selection. This is especially important considering the extremely high amputation rate for the remaining leg in diabetic patients within the first few years after the loss of the first leg. Based on the results of this study, the design of the prosthetic foot, especially the energy-storing characteristics of the device, would appear to have minimal effect on the magnitude of plantar pressures being applied to the intact limb in the insensate diabetic amputee. Thus, the prosthetist should not base the selection of the prosthetic foot on the possibility of reducing peak pressures in the remaining intact limb, but rather select the component that provides the amputee with the most optimal gait pattern and functional usage.

Conclusion

The results of this investigation indicate that the four prosthetic feet evaluated in the current study, especially those engineered to have greater energy-storing properties, do not significantly decrease the magnitude and duration of pressures acting on the plantar surface of the foot. In light of these results, prosthetic foot selection based on the possibility of reducing plantar pressures acting on the remaining limb of the unilateral diabetic amputee would not be warranted.

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