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Finite element analysis of three commonly used external fixation devices for treating Type III pilon fractures
Medical Engineering & Physics 36 (2014) 1322–1330
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Technical note
Finite element analysis of three commonly used external fixation devices for treating Type III pilon fractures Muhammad Hanif Ramlee a , Mohammed Rafiq Abdul Kadir a,∗ , Malliga Raman Murali b , Tunku Kamarul b a Medical Devices and Technology Group (MEDITEG), Faculty of Biosciences and Medical Engineering, Universiti Teknologi Malaysia, 81310 Johor Bahru, Johor, Malaysia b Tissue Engineering Group (TEG), National Orthopaedic Centre of Excellence in Research and Learning (NOCERAL), Department of Orthopaedic Surgery, Faculty ofMedicine, University of Malaya, 50603 Lembah Pantai, Kuala Lumpur, Malaysia
a r t i c l e
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Article history: Received 28 October 2013 Received in revised form 19 May 2014 Accepted 24 May 2014 Keywords: Finite element External fixator Pilon fractures Stability Biomechanics Micromovement
a b s t r a c t Pilon fractures are commonly caused by high energy trauma and can result in long-term immobilization of patients. The use of an external fixator i.e. the (1) Delta, (2) Mitkovic or (3) Unilateral frame for treating type III pilon fractures is generally recommended by many experts owing to the stability provided by these constructs. This allows this type of fracture to heal quickly whilst permitting early mobilization. However, the stability of one fixator over the other has not been previously demonstrated. This study was conducted to determine the biomechanical stability of these external fixators in type III pilon fractures using finite element modelling. Three-dimensional models of the tibia, fibula, talus, calcaneus, navicular, cuboid, three cuneiforms and five metatarsal bones were reconstructed from previously obtained CT datasets. Bones were assigned with isotropic material properties, while the cartilage was assigned as hyperelastic springs with Mooney–Rivlin properties. Axial loads of 350 N and 70 N were applied at the tibia to simulate the stance and the swing phase of a gait cycle. To prevent rigid body motion, the calcaneus and metatarsals were fixed distally in all degrees of freedom. The results indicate that the model with the Delta frame produced the lowest relative micromovement (0.03 mm) compared to the Mitkovic (0.05 mm) and Unilateral (0.42 mm) fixators during the stance phase. The highest stress concentrations were found at the pin of the Unilateral external fixator (509.2 MPa) compared to the Mitkovic (286.0 MPa) and the Delta (266.7 MPa) frames. In conclusion, the Delta external fixator was found to be the most stable external fixator for treating type III pilon fractures. © 2014 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction A pilon fracture is a general description of a comminuted fracture at the distal tibia involving the ankle joint that occurs as the result of high-energy vertical axial loading. This can occur as the result of a fall from a substantial height, road traffic accidents, industrial mishaps or sporting injuries, especially those involving contact sports [1–6]. These fractures are uncommon and represent up to 7–10% of tibia fractures and less than 1% of all lower extremity fractures [4,7]. The mechanism of injury varies from simple rotational fractures to high energy axial compression injuries complicated by shearing, rotation and bending forces [4,8,9]. In 1969,
∗ Corresponding author. Tel.: +6 07 5535961; fax: +6 07 5536222. E-mail addresses: [email protected] (M.H. Ramlee), rafi[email protected] (M.R. Abdul Kadir), [email protected] (M.R. Murali), [email protected] (T. Kamarul). http://dx.doi.org/10.1016/j.medengphy.2014.05.015 1350-4533/© 2014 IPEM. Published by Elsevier Ltd. All rights reserved.
Ruedi and Allgower [10–16] classified the pilon fractures into three types: type I is an intra-articular fracture of the distal tibia with or without minimal displacement; type II is a displaced intra-articular fracture with or without minimal comminution; and type III has significant comminution and impaction of the intra-articular surface with displacement. Treatment of pilon fractures is targeted to reduce the fracture, align the ankle position, provide fast soft tissue healing, be minimally invasive, allow the recreation of the joint surfaces and provide early ankle function [17–19]. Type I and type II fractures can be almost effortlessly restored using internal fixation, and the results are promising without any major complications [4,17,20,21]. However, the treatment of a type III fracture is still controversial since it involves an intra-articular fracture with displacement, significant comminution and is associated with high rate of complications [1,22]. Immediate treatment is reported to produce complications such as infection, loss of reduction, nonunion, malunion and deformity [23–26]. A systemic step-wise
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approach, consisting of fibular plating and temporary bridging external fixation, later substituted by a definitive external fixation, was reported to be favourable for type III fracture treatment [2,3,5]. The range of external fixators used for the treatment of type III pilon fractures includes spanning fixator with rods and clamps, dynamic or articulated devices, and ring frames [1–3,5,27–30]. The clinical outcome of using external fixators has been reported to be superior over those treated with internal plates and screws [28,29]. The three most popular spanning external fixators in the literature include the Delta, Mitkovic and Unilateral devices [7,27,30]. Although the mid-term clinical outcomes have been shown to be good in terms of successful healing process, the biomechanical stability of these devices has not been well investigated. Furthermore, there is no clear evidence in the present literature with regard to these types of implants and whether one of these will produce the best stability in treating type III pilon fractures. Therefore, the overall aim of this study was to understand the underlying biomechanics of the three most commonly used external fixators for type III pilon fractures. Finite element method was used to (1) assess the stability of the aforementioned three external fixators (2) investigate the stress distribution in the fixator and bone to highlight the likelihood of the particular areas of the implant and bone that may be subjected to excessive mechanical stress. 2. Materials and methods 2.1. Three-dimensional modelling CT images of the right lower limb used in this study were acquired with the approval of the medical ethical committee of the Hospital Tengku Ampuan Afzan, Kuantan, Malaysia [65]. The slice thickness of the CT images was 1.5 mm in a 512 × 512 matrix. The DICOM data sets, which consist of 225 CT images, were then imported into Mimics 15.1 software (Materialise, Leuven, Belgium) to reconstruct the surface geometry of the tibia, fibula, talus, calcaneus, cuboid, navicular, cuneiforms and metatarsals. The tibial bone
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was cut approximately 20 cm above the medial tibial malleolus [31,32]. A threshold of 700 Hounsfield units was used to differentiate between cortical and cancellous bone [33]. To simulate type III pilon fractures, a total of eight fragments (Fig. 1) were modelled based on the Ruedi and Allgower classification [10–16,30]. A perfect fit of the interfragments was simulated, i.e. there were no fracture gaps between the fragments. However, the fragments were allowed to move relative to each other with an assigned friction coefficient of 0.3 [34]. All contact between articulating surfaces was assigned a friction coefficient of 0.3 [34]. All three-dimensional models of the bones were converted to a surface triangular mesh and saved in a stereolithographic file format. 2.2. Cartilage and ligaments The cartilage was modelled manually with an estimated uniform thickness of 1 mm for the tibia, fibula, talus and calcaneus (Fig. 1) [35,36]. The behaviour of the cartilage was simulated using the Mooney–Rivlin hyper-elastic material properties with coefficients of C01 = 0.41 MPa and C10 = 4.1 MPa [37–40]. A total of 34 ligaments and three plantar fascias were modelled using linear spring elements with an assigned specific stiffness (Table 1). The use of linear links to simulate the ligaments was found to be adequate and has been reported in previous studies [40–43]. Multiple parallel linear springs were used to better mimic the distribution of the origin and insertion of the ligaments [39,40]. The position and insertion points of all the ligaments were estimated based on Netter (2003) [44]. 2.3. Finite element modelling The bones and cartilage in the STL files were imported into Marc.Mentat (MSC.Software, Santa Ana, CA). The software was used to convert the completed three-dimensional model to linear first order tetrahedral elements. Bones were assigned using linear isotropic material properties with elastic modulus of 7300 MPa for cortical bone [50,51] and 1100 MPa for cancellous bone [52].
Fig. 1. Finite element model; (a) Delta frame, (b) Mitkovic fixation, (c) Unilateral external fixator. Fragment 8 is located at the posterior of distal tibia.
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Table 1 Stiffness of ligaments. Ligaments represented in the models
Stiffness (N/mm)
References
Interosseous membrane (4 ligaments) Anterior tibiofibular (distal) Posterior tibiofibular (distal) Anterior talofibular Posterior talofibular Calcaneofibular Anterior tibiotalar Posterior tibiotalar Tibiocalcaneal Tibionavicular Interosseous talocalcaneal Lateral talocalcaneal Medial talocalcaneal Posterior talocalcaneal Dorsal talonavicular (2 ligaments) Calcaneonavicular (dorsal & plantar) Calcaeocuboid (dorsal & short plantar) Cuboideonavicular (dorsal & plantar) Cuneonavicular (dorsal & plantar) Intercuneiform (dorsal & plantar) Tarsometatarsal (dorsal & plantar) Metatarsal (dorsal & plantar) Medial plantar fascia Central plantar fascia Lateral plantar fascia Long plantar
400 78 101 90 70 70 70 80 122 40 70 70 70 70 70 70 70 70 70 70 70 70 200 230 180 70
[45] [46] [47] [46] [48] [48] [46] [46] [46] [46] [45,48] [48] [48] [48] [46] [46,48] [46,48] [45,46] a a a a [49] [49] [49] [45]
a: assumed from neighboring ligaments.
Poisson’s ratio for both cortical and cancellous bones was simulated with value of 0.3 [50,51] and 0.26 [52], respectively. Three external fixator frames, i.e. the Delta, Mitkovic and Unilateral systems, were designed using three-dimensional (3D) Computer Aided Design (CAD) software (Solidworks 2012, Dassault Systemes Solidworks Corp., USA) with 5 mm pins and 11 mm rods. To simulate the Delta and Mitkovic fixators, two pins were positioned at the tibial diaphysis, one pin at the body of the calcaneus and another pin at the first metatarsal (Fig. 1) [27,30]. For the Unilateral frame, only one pin was positioned at the tibial diaphysis and another pin at the body of calcaneus (Fig. 1) [7]. All the fixators were meshed using 3-Matic 7.1 (Materialise, Belgium) and were assigned with titanium material properties with a Young’s modulus of 110,000 MPa and Poisson’s ratio of 0.3 [53,54]. Mesh convergence analyses were performed and resulted in a variation in mesh size throughout the model. The smallest mesh of size 1 was used for the pin-bone interface, whereas a larger mesh size of 3 was used for the bone. The total number of elements and nodes for the Delta fixator was 675,000 and 157,000, respectively; for the Mitkovic fixator, this was 588,000 and 140,000, respectively, and for the Unilateral frame, this was 510,000 and 112,000, respectively. Contact condition between the external fixators and bone was set as an explicit contact with a tangential friction coefficient of 0.4 [52,55,56]. Radial pre-stress was not modelled at the interface between the bone and the fixators. 2.4. Boundary conditions In order to simulate human walking conditions, two physiological loads were applied in this study: (1) the swing phase [57–59] and (2) the stance phase [60,61] of a gait cycle, where the force value was determined from the adjacent muscles such as the gastrocnemius and soleus. For the swing phase, 10% of the body weight was recorded on these particular muscles [57–59]. We assumed a body weight of 70 kg in our study; therefore, 70 N was applied to the tibia in the axial direction to simulate the swing phase. For the stance phase, 50% (350 N) of the body weight was applied onto the foot, as has been reported by Cheung et al. and Simkin [60,61]. The
use of axial weight loading has become popular since this technique is a way of testing bone quality and bone healing process [62]. In order to prevent rigid body movements during the analysis, the distal surfaces of the calcaneus and all metatarsal bones were fixed in all directions (Fig. 1). The relative micromovement of all simulated models were measured between the proximal and distal fragments at the lateral side. 3. Results 3.1. Stress distribution The von Mises stress at the pin-bone interface at the tibia and calcaneus is shown in Fig. 2. During the swing phase, the observed peak values for the pin-bone interface at the tibia were 24.9 MPa, 35.6 MPa and 84.9 MPa for the Delta, Mitkovic and Unilateral fixation devices, respectively. The difference in magnitude was even higher during the stance phase for the pin-bone interface at the tibia, where the Unilateral showed two times greater stress (399.0 MPa) than the Mitkovic (206.6 MPa) and three times greater stress than the Delta fixator (130.3 MPa). Generally, at the tibia, the peak von Mises stress was found at the entrance cortex of the pin-bone interface. During the swing phase, the magnitude of the maximum stress was 0.8 MPa, 3.6 MPa and 4.5 MPa for the Delta, Mitkovic and Unilateral fixators, respectively. On the other hand, the FE results in terms of von Mises stress were greater for the simulated stance phase where the Unilateral generated at least 1.4 times greater (121.4 MPa) than the Mitkovic (87.2 MPa) and 27 times greater than the Delta (4.5 MPa) frames. At the calcaneus, high von Mises stress was observed at the pin-bone interface for both the swing and the stance phase. During the swing phase, the magnitude of stress for the Unilateral fixator was at least four times greater (98.5 MPa) as compared to the Mitkovic (20.7 MPa) and Delta (7.9 MPa) frames. Additionally, greater stresses were observed during the stance phase, where the Unilateral fixator produced 382.5 MPa, 3.9 times larger than the Mitkovic (98.0 MPa) and 9.2 times larger than Delta external fixator (41.5 MPa) The stress distribution amongst the three external fixators is illustrated in Fig. 3. During the swing phase, higher von Mises stresses were predicted at the calcaneus pin for the Unilateral external fixator (113.6 MPa) followed by the Mitkovic (80.0 MPa) and the Delta (45.6 MPa) frame. For the Delta and Mitkovic systems, the proximal pin-bone interface of the tibia bone produced a small stress of 25.0 MPa and 44.5 MPa, respectively, as compared to the distal pin-bone interface with a value of 27.0 MPa and 54.5 MPa, respectively. At the first metatarsal bone, high stress of 43.5 MPa was found for the Mitkovic fixator, whilst the Delta frame only showed 7.5 MPa of stress at that particular bone. During the stance phase, the stress at the calcaneus pin for the Unilateral external fixator (509.2 MPa) was at least 1.8 times greater than with the Mitkovic (286.0 MPa) and Delta (266.7 MPa) fixators. At the proximal pin of the Mitkovic frame, the von Mises magnitude (125.9 MPa) was close to magnitude of the Delta frame (120.2 MPa). On the other hand, a small stress of 49.5 MPa was generated at the first metatarsal pin of the Delta external fixator, whilst the Mitkovic frame (286.0 MPa) showed 4.5 times greater stress. 3.2. Displacement and micromovement Fig. 4 shows the displacement plot for the tibia bone. For the swing phase, the greatest relative micromovement was observed for the Unilateral fixator (0.3 mm) as compared to the Mitkovic and Delta (0.02 mm) frames. In contrast, the relative micromovement was higher during the stance phase. The Unilateral fixator (0.42 mm) generated 8.4 times greater micromovement than the
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Fig. 2. The von Mises stress distribution for the tibia and calcaneus bones for (a) Delta frame, (b) Mitkovic fixation and (c) Unilateral external fixation.
Mitkovic (0.05 mm) and 14 times greater micromovement than the Delta (0.03 mm) external fixators. The contour plot for the displacement of external fixators is shown in Fig. 5. When simulating the swing phase, it was demonstrated that the maximum displacement produced by the Unilateral fixator (8.7 mm) was at least three times greater than the displacement produced by the Mitkovic (3.0 mm) and Delta (0.8 mm) systems. During the stance phase, the Unilateral fixator generated the highest magnitude of displacement (34.8 mm). The Delta system was the most stable fixation system among the three constructs examined with a maximum displacement of 3.8 mm. The Mitkovic system showed a maximum displacement value of 13.4 mm.
4. Discussion The use of an external fixator for treating type III pilon fractures is a well-accepted surgical option. This system not only allows minimally invasive surgery of the soft tissue to be performed, but also maintains the ankle alignment whilst allowing early ankle mobilization [17–19]. The results of using an external fixator such as the Delta, Mitkovic or Unilateral frame have been shown to produce favourable clinical outcomes as compared to internal fixators [17–19,23–26]. Stable fixation and early anatomic reduction of all fractures and dislocations can minimize long-term morbidity and hasten soft tissue healing [63,64]. The biomechanical
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Fig. 3. The von Mises stress plot for the external fixator during swing and stance phase. (a) Delta, (b) Mitkovic and (c) Unilateral.
stability produced by these fixators, provides improved fracture healing to occur. A correlation between improved stability (albeit still allowing micromotion to occur) with improved healing has been demonstrated in a previous study [65]. However, it is clear that from our extensive literature review that a comparison of the most common constructs used for type III pilon fracture treatments appears to be lacking. Hence, the results presented in the present study are the first that we are aware of. The results presented here are of value not only for future research, but also serve as an objective measure for surgeons to justify the choice of one construct
over the other, although it may be necessary for further studies to be conducted in order to support the choices made. To build confidence in our FE results, we corroborated our data with experimental studies of Bergmann et al. [66] and Wang et al. [32]. The former measured hip contact and ground reaction forces for four patients during the most frequent activities of daily living. In order to compare our results with the work of Bergmann et al., we simulated the ankle region of interest without a fixator. Our results showed minimum and maximum ground reaction forces of 0.1 N and 175.9 N, respectively, compared to 0.1 N and 108.7 N obtained
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Fig. 4. The displacement plot for the tibia bone; (a) Delta fixation, (b) Mitkovic frame and (c) Unilateral external fixator.
by Bergmann et al. [66]. Wang et al. [32], on the other hand, used nine ankles from cadavers to measure the contact pressure of the subtalar joint using pressure-sensitive films. For the 600 N load applied at the tibia, they reported a maximum contact pressure of 5.13 ± 1.16 MPa. This is close to the result that we obtained from our FE simulation, which was 5.48 MPa.
In the present study, the high stress at the pin-bone interface and the surrounding tissues were in agreement with previous reports [67,68]. The stress concentrated at this particular region is one of the causes of an unstable external fixator construct and can lead to pain and implant loosening [12,69–72]. However, this cannot be avoided when using an external fixator since any designed implants
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Fig. 5. The displacement plot for the external fixator during the swing and stance phase. (a) Delta, (b) Mitkovic and (c) Unilateral.
using this concept will inherently have to face this issue. Nevertheless, our results showed a more favourable result for the Delta external fixator than the Mitkovic and Unilateral systems due to the lower stress magnitudes as well as better stability with lower relative micromovement values. Our findings appear to be supported by previously published clinical studies such as those by Cheema et al. [27]. In their study, they observed that the design of Delta frames provided the most stable construct, while preventing a high incidence of getting deformities [27]. For long-term clinical results, the use of the Unilateral external fixator should be avoided since the finite element predictions showed the highest relative micromovement as compared to the Delta and the Mitkovic constructs. In addition, the Unilateral system, which does not use a pin at the first metatarsal bone, can cause forefoot equinus deformities [27]. Interfragment micromovement at the fracture site has been reported as an important parameter that will assist in bone healing process. Several studies using ovine has demonstrated that micromovement between 0.15 mm and 0.4 mm can assist in the healing
of a fracture gap not more than 3 mm [73,74]. In our analysis, we simulated a perfect fit for the interfragments, so that there were no fracture gaps present. The relative micromovement for the swing phase for these bone fragments was therefore less than 0.02 mm (Mitkovic and Delta) and 0.3 mm (Unilateral). During the the stance phase, micromotions were observed to 0.03 mm (Delta), 0.05 mm (Mitkovic) and 0.42 mm (Unilateral). This suggests that in vivo, bone regeneration can be expected. Nevertheless, it is worth noting that unstable fracture fixation leads to increasing stresses on implants [75–77]. In several studies it has been shown that stress onto implants fixed at unstable sites have recoded levels in access maximum principal of 370–600 MPa [75] and von Mises stresses of 436–750 MPa [76,77] for the plates. Similarly in our analysis, all three external fixators also demonstrated maximum stress magnitudes at the fixator pin. It is fortunate however, that the magnitude of these stresses did not exceed the ultimate strength of titanium alloys used in our simulation (800–900 MPa) thus suggesting that the construct for all
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fixations appears to provide adequate stability with minimal risk of implant failure [78]. In treating type III pilon fractures, initial considerations such as pin placement and the type of external fixator must be made properly before the surgery can be conducted. Previous clinical reports have mentioned that misplacing the pins or improper use of these devices can lead to a high incidence of complications, with pin infection and loosening in up to 50% of cases and malunion rates of up to 45% [24,26,79]. In the simulations conducted here, we only attempted to show the comparison of ankle external fixator in terms of the biomechanical properties. Although the Unilateral frame showed larger displacements and relative micromovements compared to the other two constructs, stability could be better achieved by placing the proximal pin closer to the fractured segments. The literature on pin placement is fairly limited, thus further study is necessary to assess the biomechanical effects of different pin orientations. As with any study, there are limitations that need to be considered so as to not to overstate the findings. In the present study, several assumptions and simplifications were made which may have resulted in alterations to the predicted data. These limitations, which are inherently particular to computer modelling involving the reconstruction of complex joints such as the ankle, will be present and are unfortunately unavoidable. First, it should be noted that the model was simplified using isotropic material properties in this simulation, thus excluding all other important factors that may influence the prediction of fracture stability. Although more complex modelling would yield more realistic outcomes, omitting such details appears to be an acceptable practice as demonstrated by many previous studies [39,40,50,51]. Further simplifications of ligaments and plantar fascias modelling using linear links again may have diverted the results from demonstrating true, real life outcomes. The elements used to simulate the ligaments were not modified, and therefore were allowed to resist both tension and compression. Though this may not mimic the actual behaviour of ligaments, the simplified properties of a spring have been used by others with acceptable accuracy [40–43]. The insertion points of the ligaments were estimated based on a reference book by Netter (2003) [44] and confirmed by an orthopaedic specialist. This was done as there are no published reports regarding the geometrical details of all 37 ligaments modelled in our study. Considering the existence of subject-specific variation in ligament insertion points, we believe that these estimations were valid. The determination of the region of interest for the analysis is another limitation of our study. Due to constraints in computing resources, we modelled only the distal half of the tibia and fibula. Nevertheless the FE analysis still valid as similar region of interest has been used by others [31,32].
5. Conclusion The results of finite element predictions suggest that the Delta frame provides better stability and generates lower construct stresses as compared to Mitkovic and Unilateral external fixators. Our data therefore suggest that Delta fixators are superior for treating this type of fracture. However, further studies are required to validate the outcome of these simulated studies, such as those of cadaveric or clinical studies.
Ethical approval Access to the CT images of the right lower limb used in this study was granted by Dr. Zainun Bt. A. Rahman, Head of Department Diagnostic Imaging and Dr. Ghazali Ismail, Chairman of Clinical Research
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Centre, Hospital Tengku Ampuan Afzan, 25100 Kuantan, Pahang Darul Makmur, Malaysia. Acknowledgement The work has been carried out using the research grants received from eScienceFund, Ministry of Science, Technology and Innovation Malaysia, FRGS Ministry of Education Malaysia, and UTM Research University Grants. More than one of the authors of this paper was supported under University of Malaya HIR-MOHE research grant. Conflict of interest None declared. References [1] Kapukaya A, Subasi M, Arslan H, Tuzuner T. Non-reducible, open tibial plafond fractures treated with a circular external fixator (is the current classification sufficient for identifying fractures in this area). Injury 2005;36:1480–7. [2] Piper KJ, Won HY, Ellis AM. Hybrid external fixation in complex tibial plateau and plafond fractures: an Australian audit of outcomes. Injury 2005;36. [3] Marsh JL, Nepola JV, Wuest TK, Osteen D, Cox K, Oppenheim W. Unilateral external fixation until healing with the dynamic axial fixator for severe open tibial fractures. J Orthop Trauma 1991;5:341–8. [4] Bonar SK, Marsh JL. Tibial plafond fractures: changing principles of treatment. J Am Acad Orthop Surg 1994;2:297–305. [5] Kapukaya A, Subasi M, Arslan H. Management of comminuted closed tibial plafond fractures using circular external fixators. Acta Orthop Belg 2005;71:582–9. [6] Mittal R, Matthews SJ, Zavras DT, Giannoudis PV. Management of ipsilateral pilon and calcaneal fractures: a report of 2 cases. J Foot Ankle Surg 2004;43:123–30. [7] Prayson MJ, Moon BS. Stabilization of the fractured tibial plafond. Oper Tech Orthop 1999;9:216–28. [8] Borrelli JJ, Ellis E. Pilon fractures: assessment and treatment. Orthop Clin N Am 2002;33:231–45. [9] Spinosa FA. Chapter 19—classification of fractures and dislocations. Foot Ankle Radiol 2003:415–51. [10] Bartolozzi P, Lavini F. Fractures of the tibia pilon. Milan: Springer; 2004. [11] Anwar R, Tuson KWR, Khan SA. Classification and diagnosis in orthopaedic trauma. New York: Cambridge University Press; 2008. [12] Hopton BP, Harris NJ. Fractures of the foot and ankle. Surgery 2010;28:502–7. [13] Ruedi T, Matter P, Allgower M. Intra-articular fractures of the distal tibial end. Hel Chir Acta 1968;35:556–82. [14] Ruedi T. The treatment of displaced metaphyseal fractures with screws and wiring systems. Orthopedics 1989;12:55–9. [15] Ruedi T, Allgower M. The operative treatment of intra-articular fractures of the lower end of the tibia. Clin Orthop Relat Res 1979;138:105–10. [16] Ruedi T. Fractures of the lower end of the tibia into the ankle joint: results 9 years after opne reduction and internal fixation. Injury 1973;5:130–4. [17] Liporace FA, Yoon RS. An adjunct to percutaneous plate insertion to obtain optimal sagittal plane alignment in the treatment of pilon fractures. J Foot Ankle Surg 2012;51:275–7. [18] Dudko S, Kusz D, Wojciechowski P, Stoltny T. Operative treatment of ankles fractures using internal osteosynthesis by a minimal surgical approach. Foot 2004;14:185–91. [19] Mauffrey C, Vasario G, Battiston B, Lewis C, Beazley J, Seligson D. Tibial pilon fractures: a review of incidence, diagnosis, treatment, and complications. Acta Orthop Belg 2011;77:432–40. [20] Pellegrini M, Cuchacovich N, Lagos L, Henriquez H, Carcuro G, Bastias C. Minimally-invasive alternatives in the treatment of distal articular tibial fractures. Fur Sprunggelenk 2012;10:37–45. [21] Wall OR, Pinder R, Faraj AA. Ender’s nail fixation of tibial pilon fractures—a safe, minimally invasive approach for high risk patients in a small district general hospital. Inj Extra 2007;38:8. [22] Etter C, Ganz R. Long-term results of tibial plafond fractures treated with open reduction and internal fixation. Arch Orthop Trauma Surg 1991;110:277–83. [23] Picanz J. Poor results mark ORIF of tibial plafond fractures. Orthop Today 1990;10:1–2. [24] Teeny S, Wiss DA, Hathaway R, Sarmiento A. Tibial plafond fractures: errors, complications, and pitfalls in operative treatment. Orthop Trans 1990;14:265–71. [25] Rammelt S, Marti RK, Raaymakers ELFB, Grass R, Zwipp H. Joint preserving reconstruction of malunited pilon fractures. Fur Sprunggelenk 2012;10:62–72. [26] Sirkin M, Sanders R, DiPasquale T, Herscovici D. A staged protocol for soft tissue management in the treatment of complex pilon fractures. J Orthop Trauma 1999;13:78–84. [27] Cheema GS, Arora S, Sabat D, Singla J, Goel N, Maini L. The results of two-staged operative management of pilon fractures—a review of 25 cases. J Clin Orthop Trauma 2011;2:104–8.
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Technical note
Finite element analysis of three commonly used external fixation devices for treating Type III pilon fractures Muhammad Hanif Ramlee a , Mohammed Rafiq Abdul Kadir a,∗ , Malliga Raman Murali b , Tunku Kamarul b a Medical Devices and Technology Group (MEDITEG), Faculty of Biosciences and Medical Engineering, Universiti Teknologi Malaysia, 81310 Johor Bahru, Johor, Malaysia b Tissue Engineering Group (TEG), National Orthopaedic Centre of Excellence in Research and Learning (NOCERAL), Department of Orthopaedic Surgery, Faculty ofMedicine, University of Malaya, 50603 Lembah Pantai, Kuala Lumpur, Malaysia
a r t i c l e
i n f o
Article history: Received 28 October 2013 Received in revised form 19 May 2014 Accepted 24 May 2014 Keywords: Finite element External fixator Pilon fractures Stability Biomechanics Micromovement
a b s t r a c t Pilon fractures are commonly caused by high energy trauma and can result in long-term immobilization of patients. The use of an external fixator i.e. the (1) Delta, (2) Mitkovic or (3) Unilateral frame for treating type III pilon fractures is generally recommended by many experts owing to the stability provided by these constructs. This allows this type of fracture to heal quickly whilst permitting early mobilization. However, the stability of one fixator over the other has not been previously demonstrated. This study was conducted to determine the biomechanical stability of these external fixators in type III pilon fractures using finite element modelling. Three-dimensional models of the tibia, fibula, talus, calcaneus, navicular, cuboid, three cuneiforms and five metatarsal bones were reconstructed from previously obtained CT datasets. Bones were assigned with isotropic material properties, while the cartilage was assigned as hyperelastic springs with Mooney–Rivlin properties. Axial loads of 350 N and 70 N were applied at the tibia to simulate the stance and the swing phase of a gait cycle. To prevent rigid body motion, the calcaneus and metatarsals were fixed distally in all degrees of freedom. The results indicate that the model with the Delta frame produced the lowest relative micromovement (0.03 mm) compared to the Mitkovic (0.05 mm) and Unilateral (0.42 mm) fixators during the stance phase. The highest stress concentrations were found at the pin of the Unilateral external fixator (509.2 MPa) compared to the Mitkovic (286.0 MPa) and the Delta (266.7 MPa) frames. In conclusion, the Delta external fixator was found to be the most stable external fixator for treating type III pilon fractures. © 2014 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction A pilon fracture is a general description of a comminuted fracture at the distal tibia involving the ankle joint that occurs as the result of high-energy vertical axial loading. This can occur as the result of a fall from a substantial height, road traffic accidents, industrial mishaps or sporting injuries, especially those involving contact sports [1–6]. These fractures are uncommon and represent up to 7–10% of tibia fractures and less than 1% of all lower extremity fractures [4,7]. The mechanism of injury varies from simple rotational fractures to high energy axial compression injuries complicated by shearing, rotation and bending forces [4,8,9]. In 1969,
∗ Corresponding author. Tel.: +6 07 5535961; fax: +6 07 5536222. E-mail addresses: [email protected] (M.H. Ramlee), rafi[email protected] (M.R. Abdul Kadir), [email protected] (M.R. Murali), [email protected] (T. Kamarul). http://dx.doi.org/10.1016/j.medengphy.2014.05.015 1350-4533/© 2014 IPEM. Published by Elsevier Ltd. All rights reserved.
Ruedi and Allgower [10–16] classified the pilon fractures into three types: type I is an intra-articular fracture of the distal tibia with or without minimal displacement; type II is a displaced intra-articular fracture with or without minimal comminution; and type III has significant comminution and impaction of the intra-articular surface with displacement. Treatment of pilon fractures is targeted to reduce the fracture, align the ankle position, provide fast soft tissue healing, be minimally invasive, allow the recreation of the joint surfaces and provide early ankle function [17–19]. Type I and type II fractures can be almost effortlessly restored using internal fixation, and the results are promising without any major complications [4,17,20,21]. However, the treatment of a type III fracture is still controversial since it involves an intra-articular fracture with displacement, significant comminution and is associated with high rate of complications [1,22]. Immediate treatment is reported to produce complications such as infection, loss of reduction, nonunion, malunion and deformity [23–26]. A systemic step-wise
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approach, consisting of fibular plating and temporary bridging external fixation, later substituted by a definitive external fixation, was reported to be favourable for type III fracture treatment [2,3,5]. The range of external fixators used for the treatment of type III pilon fractures includes spanning fixator with rods and clamps, dynamic or articulated devices, and ring frames [1–3,5,27–30]. The clinical outcome of using external fixators has been reported to be superior over those treated with internal plates and screws [28,29]. The three most popular spanning external fixators in the literature include the Delta, Mitkovic and Unilateral devices [7,27,30]. Although the mid-term clinical outcomes have been shown to be good in terms of successful healing process, the biomechanical stability of these devices has not been well investigated. Furthermore, there is no clear evidence in the present literature with regard to these types of implants and whether one of these will produce the best stability in treating type III pilon fractures. Therefore, the overall aim of this study was to understand the underlying biomechanics of the three most commonly used external fixators for type III pilon fractures. Finite element method was used to (1) assess the stability of the aforementioned three external fixators (2) investigate the stress distribution in the fixator and bone to highlight the likelihood of the particular areas of the implant and bone that may be subjected to excessive mechanical stress. 2. Materials and methods 2.1. Three-dimensional modelling CT images of the right lower limb used in this study were acquired with the approval of the medical ethical committee of the Hospital Tengku Ampuan Afzan, Kuantan, Malaysia [65]. The slice thickness of the CT images was 1.5 mm in a 512 × 512 matrix. The DICOM data sets, which consist of 225 CT images, were then imported into Mimics 15.1 software (Materialise, Leuven, Belgium) to reconstruct the surface geometry of the tibia, fibula, talus, calcaneus, cuboid, navicular, cuneiforms and metatarsals. The tibial bone
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was cut approximately 20 cm above the medial tibial malleolus [31,32]. A threshold of 700 Hounsfield units was used to differentiate between cortical and cancellous bone [33]. To simulate type III pilon fractures, a total of eight fragments (Fig. 1) were modelled based on the Ruedi and Allgower classification [10–16,30]. A perfect fit of the interfragments was simulated, i.e. there were no fracture gaps between the fragments. However, the fragments were allowed to move relative to each other with an assigned friction coefficient of 0.3 [34]. All contact between articulating surfaces was assigned a friction coefficient of 0.3 [34]. All three-dimensional models of the bones were converted to a surface triangular mesh and saved in a stereolithographic file format. 2.2. Cartilage and ligaments The cartilage was modelled manually with an estimated uniform thickness of 1 mm for the tibia, fibula, talus and calcaneus (Fig. 1) [35,36]. The behaviour of the cartilage was simulated using the Mooney–Rivlin hyper-elastic material properties with coefficients of C01 = 0.41 MPa and C10 = 4.1 MPa [37–40]. A total of 34 ligaments and three plantar fascias were modelled using linear spring elements with an assigned specific stiffness (Table 1). The use of linear links to simulate the ligaments was found to be adequate and has been reported in previous studies [40–43]. Multiple parallel linear springs were used to better mimic the distribution of the origin and insertion of the ligaments [39,40]. The position and insertion points of all the ligaments were estimated based on Netter (2003) [44]. 2.3. Finite element modelling The bones and cartilage in the STL files were imported into Marc.Mentat (MSC.Software, Santa Ana, CA). The software was used to convert the completed three-dimensional model to linear first order tetrahedral elements. Bones were assigned using linear isotropic material properties with elastic modulus of 7300 MPa for cortical bone [50,51] and 1100 MPa for cancellous bone [52].
Fig. 1. Finite element model; (a) Delta frame, (b) Mitkovic fixation, (c) Unilateral external fixator. Fragment 8 is located at the posterior of distal tibia.
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Table 1 Stiffness of ligaments. Ligaments represented in the models
Stiffness (N/mm)
References
Interosseous membrane (4 ligaments) Anterior tibiofibular (distal) Posterior tibiofibular (distal) Anterior talofibular Posterior talofibular Calcaneofibular Anterior tibiotalar Posterior tibiotalar Tibiocalcaneal Tibionavicular Interosseous talocalcaneal Lateral talocalcaneal Medial talocalcaneal Posterior talocalcaneal Dorsal talonavicular (2 ligaments) Calcaneonavicular (dorsal & plantar) Calcaeocuboid (dorsal & short plantar) Cuboideonavicular (dorsal & plantar) Cuneonavicular (dorsal & plantar) Intercuneiform (dorsal & plantar) Tarsometatarsal (dorsal & plantar) Metatarsal (dorsal & plantar) Medial plantar fascia Central plantar fascia Lateral plantar fascia Long plantar
400 78 101 90 70 70 70 80 122 40 70 70 70 70 70 70 70 70 70 70 70 70 200 230 180 70
[45] [46] [47] [46] [48] [48] [46] [46] [46] [46] [45,48] [48] [48] [48] [46] [46,48] [46,48] [45,46] a a a a [49] [49] [49] [45]
a: assumed from neighboring ligaments.
Poisson’s ratio for both cortical and cancellous bones was simulated with value of 0.3 [50,51] and 0.26 [52], respectively. Three external fixator frames, i.e. the Delta, Mitkovic and Unilateral systems, were designed using three-dimensional (3D) Computer Aided Design (CAD) software (Solidworks 2012, Dassault Systemes Solidworks Corp., USA) with 5 mm pins and 11 mm rods. To simulate the Delta and Mitkovic fixators, two pins were positioned at the tibial diaphysis, one pin at the body of the calcaneus and another pin at the first metatarsal (Fig. 1) [27,30]. For the Unilateral frame, only one pin was positioned at the tibial diaphysis and another pin at the body of calcaneus (Fig. 1) [7]. All the fixators were meshed using 3-Matic 7.1 (Materialise, Belgium) and were assigned with titanium material properties with a Young’s modulus of 110,000 MPa and Poisson’s ratio of 0.3 [53,54]. Mesh convergence analyses were performed and resulted in a variation in mesh size throughout the model. The smallest mesh of size 1 was used for the pin-bone interface, whereas a larger mesh size of 3 was used for the bone. The total number of elements and nodes for the Delta fixator was 675,000 and 157,000, respectively; for the Mitkovic fixator, this was 588,000 and 140,000, respectively, and for the Unilateral frame, this was 510,000 and 112,000, respectively. Contact condition between the external fixators and bone was set as an explicit contact with a tangential friction coefficient of 0.4 [52,55,56]. Radial pre-stress was not modelled at the interface between the bone and the fixators. 2.4. Boundary conditions In order to simulate human walking conditions, two physiological loads were applied in this study: (1) the swing phase [57–59] and (2) the stance phase [60,61] of a gait cycle, where the force value was determined from the adjacent muscles such as the gastrocnemius and soleus. For the swing phase, 10% of the body weight was recorded on these particular muscles [57–59]. We assumed a body weight of 70 kg in our study; therefore, 70 N was applied to the tibia in the axial direction to simulate the swing phase. For the stance phase, 50% (350 N) of the body weight was applied onto the foot, as has been reported by Cheung et al. and Simkin [60,61]. The
use of axial weight loading has become popular since this technique is a way of testing bone quality and bone healing process [62]. In order to prevent rigid body movements during the analysis, the distal surfaces of the calcaneus and all metatarsal bones were fixed in all directions (Fig. 1). The relative micromovement of all simulated models were measured between the proximal and distal fragments at the lateral side. 3. Results 3.1. Stress distribution The von Mises stress at the pin-bone interface at the tibia and calcaneus is shown in Fig. 2. During the swing phase, the observed peak values for the pin-bone interface at the tibia were 24.9 MPa, 35.6 MPa and 84.9 MPa for the Delta, Mitkovic and Unilateral fixation devices, respectively. The difference in magnitude was even higher during the stance phase for the pin-bone interface at the tibia, where the Unilateral showed two times greater stress (399.0 MPa) than the Mitkovic (206.6 MPa) and three times greater stress than the Delta fixator (130.3 MPa). Generally, at the tibia, the peak von Mises stress was found at the entrance cortex of the pin-bone interface. During the swing phase, the magnitude of the maximum stress was 0.8 MPa, 3.6 MPa and 4.5 MPa for the Delta, Mitkovic and Unilateral fixators, respectively. On the other hand, the FE results in terms of von Mises stress were greater for the simulated stance phase where the Unilateral generated at least 1.4 times greater (121.4 MPa) than the Mitkovic (87.2 MPa) and 27 times greater than the Delta (4.5 MPa) frames. At the calcaneus, high von Mises stress was observed at the pin-bone interface for both the swing and the stance phase. During the swing phase, the magnitude of stress for the Unilateral fixator was at least four times greater (98.5 MPa) as compared to the Mitkovic (20.7 MPa) and Delta (7.9 MPa) frames. Additionally, greater stresses were observed during the stance phase, where the Unilateral fixator produced 382.5 MPa, 3.9 times larger than the Mitkovic (98.0 MPa) and 9.2 times larger than Delta external fixator (41.5 MPa) The stress distribution amongst the three external fixators is illustrated in Fig. 3. During the swing phase, higher von Mises stresses were predicted at the calcaneus pin for the Unilateral external fixator (113.6 MPa) followed by the Mitkovic (80.0 MPa) and the Delta (45.6 MPa) frame. For the Delta and Mitkovic systems, the proximal pin-bone interface of the tibia bone produced a small stress of 25.0 MPa and 44.5 MPa, respectively, as compared to the distal pin-bone interface with a value of 27.0 MPa and 54.5 MPa, respectively. At the first metatarsal bone, high stress of 43.5 MPa was found for the Mitkovic fixator, whilst the Delta frame only showed 7.5 MPa of stress at that particular bone. During the stance phase, the stress at the calcaneus pin for the Unilateral external fixator (509.2 MPa) was at least 1.8 times greater than with the Mitkovic (286.0 MPa) and Delta (266.7 MPa) fixators. At the proximal pin of the Mitkovic frame, the von Mises magnitude (125.9 MPa) was close to magnitude of the Delta frame (120.2 MPa). On the other hand, a small stress of 49.5 MPa was generated at the first metatarsal pin of the Delta external fixator, whilst the Mitkovic frame (286.0 MPa) showed 4.5 times greater stress. 3.2. Displacement and micromovement Fig. 4 shows the displacement plot for the tibia bone. For the swing phase, the greatest relative micromovement was observed for the Unilateral fixator (0.3 mm) as compared to the Mitkovic and Delta (0.02 mm) frames. In contrast, the relative micromovement was higher during the stance phase. The Unilateral fixator (0.42 mm) generated 8.4 times greater micromovement than the
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Fig. 2. The von Mises stress distribution for the tibia and calcaneus bones for (a) Delta frame, (b) Mitkovic fixation and (c) Unilateral external fixation.
Mitkovic (0.05 mm) and 14 times greater micromovement than the Delta (0.03 mm) external fixators. The contour plot for the displacement of external fixators is shown in Fig. 5. When simulating the swing phase, it was demonstrated that the maximum displacement produced by the Unilateral fixator (8.7 mm) was at least three times greater than the displacement produced by the Mitkovic (3.0 mm) and Delta (0.8 mm) systems. During the stance phase, the Unilateral fixator generated the highest magnitude of displacement (34.8 mm). The Delta system was the most stable fixation system among the three constructs examined with a maximum displacement of 3.8 mm. The Mitkovic system showed a maximum displacement value of 13.4 mm.
4. Discussion The use of an external fixator for treating type III pilon fractures is a well-accepted surgical option. This system not only allows minimally invasive surgery of the soft tissue to be performed, but also maintains the ankle alignment whilst allowing early ankle mobilization [17–19]. The results of using an external fixator such as the Delta, Mitkovic or Unilateral frame have been shown to produce favourable clinical outcomes as compared to internal fixators [17–19,23–26]. Stable fixation and early anatomic reduction of all fractures and dislocations can minimize long-term morbidity and hasten soft tissue healing [63,64]. The biomechanical
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Fig. 3. The von Mises stress plot for the external fixator during swing and stance phase. (a) Delta, (b) Mitkovic and (c) Unilateral.
stability produced by these fixators, provides improved fracture healing to occur. A correlation between improved stability (albeit still allowing micromotion to occur) with improved healing has been demonstrated in a previous study [65]. However, it is clear that from our extensive literature review that a comparison of the most common constructs used for type III pilon fracture treatments appears to be lacking. Hence, the results presented in the present study are the first that we are aware of. The results presented here are of value not only for future research, but also serve as an objective measure for surgeons to justify the choice of one construct
over the other, although it may be necessary for further studies to be conducted in order to support the choices made. To build confidence in our FE results, we corroborated our data with experimental studies of Bergmann et al. [66] and Wang et al. [32]. The former measured hip contact and ground reaction forces for four patients during the most frequent activities of daily living. In order to compare our results with the work of Bergmann et al., we simulated the ankle region of interest without a fixator. Our results showed minimum and maximum ground reaction forces of 0.1 N and 175.9 N, respectively, compared to 0.1 N and 108.7 N obtained
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Fig. 4. The displacement plot for the tibia bone; (a) Delta fixation, (b) Mitkovic frame and (c) Unilateral external fixator.
by Bergmann et al. [66]. Wang et al. [32], on the other hand, used nine ankles from cadavers to measure the contact pressure of the subtalar joint using pressure-sensitive films. For the 600 N load applied at the tibia, they reported a maximum contact pressure of 5.13 ± 1.16 MPa. This is close to the result that we obtained from our FE simulation, which was 5.48 MPa.
In the present study, the high stress at the pin-bone interface and the surrounding tissues were in agreement with previous reports [67,68]. The stress concentrated at this particular region is one of the causes of an unstable external fixator construct and can lead to pain and implant loosening [12,69–72]. However, this cannot be avoided when using an external fixator since any designed implants
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Fig. 5. The displacement plot for the external fixator during the swing and stance phase. (a) Delta, (b) Mitkovic and (c) Unilateral.
using this concept will inherently have to face this issue. Nevertheless, our results showed a more favourable result for the Delta external fixator than the Mitkovic and Unilateral systems due to the lower stress magnitudes as well as better stability with lower relative micromovement values. Our findings appear to be supported by previously published clinical studies such as those by Cheema et al. [27]. In their study, they observed that the design of Delta frames provided the most stable construct, while preventing a high incidence of getting deformities [27]. For long-term clinical results, the use of the Unilateral external fixator should be avoided since the finite element predictions showed the highest relative micromovement as compared to the Delta and the Mitkovic constructs. In addition, the Unilateral system, which does not use a pin at the first metatarsal bone, can cause forefoot equinus deformities [27]. Interfragment micromovement at the fracture site has been reported as an important parameter that will assist in bone healing process. Several studies using ovine has demonstrated that micromovement between 0.15 mm and 0.4 mm can assist in the healing
of a fracture gap not more than 3 mm [73,74]. In our analysis, we simulated a perfect fit for the interfragments, so that there were no fracture gaps present. The relative micromovement for the swing phase for these bone fragments was therefore less than 0.02 mm (Mitkovic and Delta) and 0.3 mm (Unilateral). During the the stance phase, micromotions were observed to 0.03 mm (Delta), 0.05 mm (Mitkovic) and 0.42 mm (Unilateral). This suggests that in vivo, bone regeneration can be expected. Nevertheless, it is worth noting that unstable fracture fixation leads to increasing stresses on implants [75–77]. In several studies it has been shown that stress onto implants fixed at unstable sites have recoded levels in access maximum principal of 370–600 MPa [75] and von Mises stresses of 436–750 MPa [76,77] for the plates. Similarly in our analysis, all three external fixators also demonstrated maximum stress magnitudes at the fixator pin. It is fortunate however, that the magnitude of these stresses did not exceed the ultimate strength of titanium alloys used in our simulation (800–900 MPa) thus suggesting that the construct for all
M.H. Ramlee et al. / Medical Engineering & Physics 36 (2014) 1322–1330
fixations appears to provide adequate stability with minimal risk of implant failure [78]. In treating type III pilon fractures, initial considerations such as pin placement and the type of external fixator must be made properly before the surgery can be conducted. Previous clinical reports have mentioned that misplacing the pins or improper use of these devices can lead to a high incidence of complications, with pin infection and loosening in up to 50% of cases and malunion rates of up to 45% [24,26,79]. In the simulations conducted here, we only attempted to show the comparison of ankle external fixator in terms of the biomechanical properties. Although the Unilateral frame showed larger displacements and relative micromovements compared to the other two constructs, stability could be better achieved by placing the proximal pin closer to the fractured segments. The literature on pin placement is fairly limited, thus further study is necessary to assess the biomechanical effects of different pin orientations. As with any study, there are limitations that need to be considered so as to not to overstate the findings. In the present study, several assumptions and simplifications were made which may have resulted in alterations to the predicted data. These limitations, which are inherently particular to computer modelling involving the reconstruction of complex joints such as the ankle, will be present and are unfortunately unavoidable. First, it should be noted that the model was simplified using isotropic material properties in this simulation, thus excluding all other important factors that may influence the prediction of fracture stability. Although more complex modelling would yield more realistic outcomes, omitting such details appears to be an acceptable practice as demonstrated by many previous studies [39,40,50,51]. Further simplifications of ligaments and plantar fascias modelling using linear links again may have diverted the results from demonstrating true, real life outcomes. The elements used to simulate the ligaments were not modified, and therefore were allowed to resist both tension and compression. Though this may not mimic the actual behaviour of ligaments, the simplified properties of a spring have been used by others with acceptable accuracy [40–43]. The insertion points of the ligaments were estimated based on a reference book by Netter (2003) [44] and confirmed by an orthopaedic specialist. This was done as there are no published reports regarding the geometrical details of all 37 ligaments modelled in our study. Considering the existence of subject-specific variation in ligament insertion points, we believe that these estimations were valid. The determination of the region of interest for the analysis is another limitation of our study. Due to constraints in computing resources, we modelled only the distal half of the tibia and fibula. Nevertheless the FE analysis still valid as similar region of interest has been used by others [31,32].
5. Conclusion The results of finite element predictions suggest that the Delta frame provides better stability and generates lower construct stresses as compared to Mitkovic and Unilateral external fixators. Our data therefore suggest that Delta fixators are superior for treating this type of fracture. However, further studies are required to validate the outcome of these simulated studies, such as those of cadaveric or clinical studies.
Ethical approval Access to the CT images of the right lower limb used in this study was granted by Dr. Zainun Bt. A. Rahman, Head of Department Diagnostic Imaging and Dr. Ghazali Ismail, Chairman of Clinical Research
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Centre, Hospital Tengku Ampuan Afzan, 25100 Kuantan, Pahang Darul Makmur, Malaysia. Acknowledgement The work has been carried out using the research grants received from eScienceFund, Ministry of Science, Technology and Innovation Malaysia, FRGS Ministry of Education Malaysia, and UTM Research University Grants. More than one of the authors of this paper was supported under University of Malaya HIR-MOHE research grant. Conflict of interest None declared. References [1] Kapukaya A, Subasi M, Arslan H, Tuzuner T. Non-reducible, open tibial plafond fractures treated with a circular external fixator (is the current classification sufficient for identifying fractures in this area). Injury 2005;36:1480–7. [2] Piper KJ, Won HY, Ellis AM. Hybrid external fixation in complex tibial plateau and plafond fractures: an Australian audit of outcomes. Injury 2005;36. [3] Marsh JL, Nepola JV, Wuest TK, Osteen D, Cox K, Oppenheim W. Unilateral external fixation until healing with the dynamic axial fixator for severe open tibial fractures. J Orthop Trauma 1991;5:341–8. [4] Bonar SK, Marsh JL. Tibial plafond fractures: changing principles of treatment. J Am Acad Orthop Surg 1994;2:297–305. [5] Kapukaya A, Subasi M, Arslan H. Management of comminuted closed tibial plafond fractures using circular external fixators. Acta Orthop Belg 2005;71:582–9. [6] Mittal R, Matthews SJ, Zavras DT, Giannoudis PV. Management of ipsilateral pilon and calcaneal fractures: a report of 2 cases. J Foot Ankle Surg 2004;43:123–30. [7] Prayson MJ, Moon BS. Stabilization of the fractured tibial plafond. Oper Tech Orthop 1999;9:216–28. [8] Borrelli JJ, Ellis E. Pilon fractures: assessment and treatment. Orthop Clin N Am 2002;33:231–45. [9] Spinosa FA. Chapter 19—classification of fractures and dislocations. Foot Ankle Radiol 2003:415–51. [10] Bartolozzi P, Lavini F. Fractures of the tibia pilon. Milan: Springer; 2004. [11] Anwar R, Tuson KWR, Khan SA. Classification and diagnosis in orthopaedic trauma. New York: Cambridge University Press; 2008. [12] Hopton BP, Harris NJ. Fractures of the foot and ankle. Surgery 2010;28:502–7. [13] Ruedi T, Matter P, Allgower M. Intra-articular fractures of the distal tibial end. Hel Chir Acta 1968;35:556–82. [14] Ruedi T. The treatment of displaced metaphyseal fractures with screws and wiring systems. Orthopedics 1989;12:55–9. [15] Ruedi T, Allgower M. The operative treatment of intra-articular fractures of the lower end of the tibia. Clin Orthop Relat Res 1979;138:105–10. [16] Ruedi T. Fractures of the lower end of the tibia into the ankle joint: results 9 years after opne reduction and internal fixation. Injury 1973;5:130–4. [17] Liporace FA, Yoon RS. An adjunct to percutaneous plate insertion to obtain optimal sagittal plane alignment in the treatment of pilon fractures. J Foot Ankle Surg 2012;51:275–7. [18] Dudko S, Kusz D, Wojciechowski P, Stoltny T. Operative treatment of ankles fractures using internal osteosynthesis by a minimal surgical approach. Foot 2004;14:185–91. [19] Mauffrey C, Vasario G, Battiston B, Lewis C, Beazley J, Seligson D. Tibial pilon fractures: a review of incidence, diagnosis, treatment, and complications. Acta Orthop Belg 2011;77:432–40. [20] Pellegrini M, Cuchacovich N, Lagos L, Henriquez H, Carcuro G, Bastias C. Minimally-invasive alternatives in the treatment of distal articular tibial fractures. Fur Sprunggelenk 2012;10:37–45. [21] Wall OR, Pinder R, Faraj AA. Ender’s nail fixation of tibial pilon fractures—a safe, minimally invasive approach for high risk patients in a small district general hospital. Inj Extra 2007;38:8. [22] Etter C, Ganz R. Long-term results of tibial plafond fractures treated with open reduction and internal fixation. Arch Orthop Trauma Surg 1991;110:277–83. [23] Picanz J. Poor results mark ORIF of tibial plafond fractures. Orthop Today 1990;10:1–2. [24] Teeny S, Wiss DA, Hathaway R, Sarmiento A. Tibial plafond fractures: errors, complications, and pitfalls in operative treatment. Orthop Trans 1990;14:265–71. [25] Rammelt S, Marti RK, Raaymakers ELFB, Grass R, Zwipp H. Joint preserving reconstruction of malunited pilon fractures. Fur Sprunggelenk 2012;10:62–72. [26] Sirkin M, Sanders R, DiPasquale T, Herscovici D. A staged protocol for soft tissue management in the treatment of complex pilon fractures. J Orthop Trauma 1999;13:78–84. [27] Cheema GS, Arora S, Sabat D, Singla J, Goel N, Maini L. The results of two-staged operative management of pilon fractures—a review of 25 cases. J Clin Orthop Trauma 2011;2:104–8.
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