Osteoarthritis of the centrodistal and tarsometatarsal joints can be a debilitating cause of lameness in many equine athletes.1,2 Pain associated with osteoarthritic changes originates in the synovial membrane, joint capsule, periarticular muscles and ligaments, periosteum, and subchondral bone.3 Instability of the joints has been associated with the initiation and propagation of pain for humans with naturally occurring osteoarthritis and horses with experimentally induced osteoarthritis.4,5 The chronic repetitive compression, torsion, and shear forces exerted can initiate and perpetuate the process of osteoarthritis in low-motion joints.6
Medical treatments for horses with osteoarthritis are designed to manage the disease. These include oral administration of NSAIDs with or without intra-articular administration of corticosteroid drugs, which may permit the continued training and performance of affected horses. Results of medical treatment are often disappointing. Horses may become refractory to treatment, with lameness persisting for 25% to 50% of them.7,8 Ankylosis of low-motion joints can develop naturally with time owing to cartilage degeneration, which permits bone bridging. However, without some form of veterinary intervention, natural ankylosis can take many years to develop and rarely results in complete bony union and return to soundness.9
In horses with osteoarthritis of low-motion joints, facilitated ankylosis and surgical arthrodesis can be performed. Techniques to promote joint fusion have been developed, but there is no reported superior method of fusion.10–15 Current strategies aimed at facilitating ankylosis include surgical drilling,10 intra-articular injection of sodium monoiodoacetate,12 intra-articular injection of ethyl alchohol,13 and laser-facilitated ankylosis.12,15 Results achieved with ethyl alcohol appear encouraging, with 88% of treated horses having radiographic fusion within 12 months after treatment, with minimal postinjection lameness caused by neurolytic properties of the injection. These methods do not increase joint stability and therefore may result in prolonged periods of pain. Horses that fail to respond to these techniques may require more aggressive intervention through internal fixation techniques to achieve soundness.
The surgical placement of implants to encourage bony fusion and increase joint stability has been used successfully for fusion of joints of the distal aspect of the tarsus, including use of T-plates and straight, narrow 4.5-mm dynamic compression plates.14,16 These techniques typically involve a large incision around the tarsus, which is complicated by the sparse soft tissue covering. Perforated stainless steel cylinders (Bagby baskets) have been used to perform arthrodesis involving the cervical vertebral bodies,17 metacarpophalangeal joint,18 and centrodistal and tarsometatarsal joints.19 Kerf-cut cylinders have been introduced as an improvement on the Bagby basket design and have similar stiffness and moment to failure as locking compression plates applied to equine cervical vertebrae.20
The purpose of the study reported here was to evaluate in vitro the biomechanical properties of 4 techniques that are used to facilitate fusion of the centrodistal and tarsometatarsal joints of horses: surgical drilling, DMLCP, kerf-cut cylinders inserted in a mediolateral direction through both centrodistal and tarsometatarsal joints (MLKC), and a DKC that spans both the centrodistal and tarsometatarsal joints. Our aim included evaluation of the 4 fixation techniques with respect to dorsoplantar bending, lateromedial bending, mediolateral bending, axial compression, and torsional loading modes. The hypotheses were that surgical drilling would result in destabilization of the centrodistal and tarsometatarsal joints, compared with properties of control joints, across the loading modes; DMLCPs would increase stability of the centrodistal and tarsometatarsal joints, compared with properties of control joints, across the loading modes; DKC involving 1 kerf-cut cylinder would be superior to MLKC involving 2 cylinders across the loading modes; and the DKC approach would be superior to the DMLCP approach across the loading modes.
Materials and Methods
Collection and preparation of specimens
Paired tarsi were obtained from cadavers of 24 adult horses < 13 years of age that had been euthanized for reasons unrelated to tarsal injury. Tarsi were radiographed prior to use in the study; those with radiographic evidence of substantial osteoarthritic changes were excluded; only tarsi with absent or minor evidence of centrodistal or tarsometatarsal osteoarthritis as identified on 4 radiographic views were accepted into the study. This mild grade included subtle changes that involved only a minor proportion of the total radiographically visible joint spaces or a single radiographic abnormality.10
Specimens were sectioned at the mid tibia and mid metatarsal region and trimmed of skin and musculature, taking care to preserve ligamentous, collateral, and capsular attachments. Specimens were then wrapped in gauze soaked in saline (0.9% NaCl) solution and frozen at −20°C until preparation for implant placement and testing. Frozen specimens were thawed overnight at room temperature (24°C) prior to implant placement.
Experimental design
Calibrated radiographic determination of tibial width was calculated for each pair of tarsi. A randomized block design, blocking on tibial width (measured on a straight dorsoplantar radiograph at the level of the distal tibial physeal scar), was used to assign tarsi pairs into 4 test groups to minimize variations in tarsal size among groups. The 2 blocking groups used were tibiae with a width ≥ 9.0 cm and those with a width < 9.0 cm. From these 2 blocking groups, tarsi were equally distributed among the 4 test groups to yield 6 pairs of tarsi/test group. Pairs of tarsi were matched and allocated equally so that the left tarsus and right tarsus were the intact control tarsus and surgical test construct in an equal number of cases. The order of the nondestructive testing modes was randomly assigned to each pair by means of a Latin square design, with axial compression consistently tested last. From preliminary data, we determined that 2 loading cycles would be required to precondition each tarsus to allow accurate data acquisition. Therefore, each testing mode was assessed through 5 loading cycles; only data from the final 3 loading cycles were used to calculate stiffness.
Test construct creation
For creation of surgical drilling constructs, drilling into the tarsus specimens was performed with a 4.5-mm drill bit. Three divergent drill tracts were created across each joint level as previously described.10 Placement of an 18-gauge needle into the joint spaces and radiographic guidance were used to correctly position the drill tracts.
For creation of MLKC constructs, mediolateral placement of kerf-cut cylinders was performed on the medial aspect of the tarsus specimen, centered in the region over the third tarsal bone immediately caudal to the location of the saphenous vein. Two 1-mm Kirschner wires were used as guides for MLKC placement: one positioned in the centrodistal joint and the other in the tarsometatarsal joint; their positions were radiographically confirmed. A 9-mm cannulated drill bit was placed over each Kirschner wire and drilled across the joint to a depth of 50 mm. A 12-mm custom-designed kerf tap was used to prepare the implant sites. Finally, 2 stainless steel kerf-cut cylinders with a 10-mm core diameter and 12-mm external threaded diametera were inserted and tightened across the joints (Figure 1).
Photograph of a kerf-cut cylinder measuring 15 mm in length and 22 mm in external threaded diameter (A) and one measuring 35 mm in length and 12 mm in external threaded diameter (B). Note that images are not of the same scale.
Citation: American Journal of Veterinary Research 77, 10; 10.2460/ajvr.77.10.1071
For creation of DKC constructs, a single stainless steel DKC measuring 15 mm in length and 22 mm in external threaded diameter was placed on the dorsomedial aspect of each tarsus specimen, centered over the third tarsal bone. A 1-mm Kirschner wire was placed in the central region of the third tarsal bone. Once the Kirschner wire was centered in a position parallel to the articular surface of the centrodistal and tarsometatarsal joints, a 13-mm cannulated drill bit was used to drill a core 15 mm into the bone. Next, a 13-mm-diameter metallic plug was inserted into the 15-mm hole to act as a guide for a 16-mm circular core cutter, which was placed over the plug and drilled into the bone 20 mm from the surface. Following this, a 16-mm-diameter metallic plug was inserted, which acted as the stabilizer for a kerf cleaner that was used to prepare the hole, followed by tapping of the hole with a modified kerf tap. Finally, the 15-mm-long implantb was firmly secured into the joint space (Figure 1).
For creation of the DMLCP constructs, tarsus specimens were surgically drilled as indicated for the surgical drilling constructs, followed by plate fixation, to simulate the processes that would occur at the time of surgery. A dorsomedial approach was used for plate placement. A locking compression T-platec was applied so the proximal 2 holes were placed in the central tarsal bone (24-mm-long, 5.0-mm-diameter locking screws), a single hole was placed in the third tarsal bone (28-mm-long, 4.5-mm-diamter cortical screw), and the distal 3 holes were placed in MT3 (60-, 48-, and 44-mm-long, 5.0-mm-diameter locking screws, proximal to distal, respectively; Figure 2).
Photograph of DMLCP placement in a tarsus specimen from an equine cadaver in relation to the 3 joint levels in the distal aspect, denoted by the white curvilinear lines (A), and radiographic image of a DMLCP construct and screw placement in a similar specimen (B).
Citation: American Journal of Veterinary Research 77, 10; 10.2460/ajvr.77.10.1071
Specimen mounting
All specimens were trimmed of bone at the proximal and distal ends to ensure a uniform length across all construct groups and to generate a consistent size and span of approximately 6.4 cm across the distal aspect of the tarsal joints for testing. Four 10-mm-diameter holes were drilled at circumferentially divergent angles in the distal aspect of the metatarsus, followed by placement of 10-mm-diameter, 14-cm-long stainless steel alloy rods through each hole to permit fixation of the distal portion. In the proximal region, a 5.5-mm cortical screw was placed across the tuber calcanei and the distal aspect of the tibia in lag fashion. A 6-hole, 4.5-mm narrow dynamic compression plate was used on the lateral aspect of the calcaneus and tibia, and 4.5-mm cortical screws were used to lag the plate to the bone. Two 10-mm holes were drilled in a mediolateral direction: one across the talus and the other across the calcaneus. Two 10-mm holes were drilled at divergent angles in the proximal aspect of the tibia, followed by the insertion of 10-mm stainless steel alloy rods through each hole to permit fixation of the proximal portion. The goal of this fixation technique was to immobilize as far as possible any movement in the tarsocrural and talocalcaneal joints (Figure 3).
Lateral radiographic view of an equine tarsus specimen in which 10-mm steel rods have been placed through various components of the joint. A dynamic compression plate was placed between the talus and the tibia, and a 5.5-mm cortical screw was used to lag the calcaneus to the tibia. The proximal portion of the construct is encased in thermal-setting plastic compound.
Citation: American Journal of Veterinary Research 77, 10; 10.2460/ajvr.77.10.1071
Once each construct was prepared, the distal portion was positioned in a 15.2 × 15.2 × 15.2-cm container. A thermal-setting plastic compoundd was then poured around the construct. After the compound had hardened, the construct was transposed vertically and the proximal portion was encased in the plastic compound. After complete construct generation, constructs were labeled and placed in cold storage (–4°C) until testing. All implant placement and specimen mounting procedures were performed by the same person (AHB).
Mechanical testing
All constructs were thawed at room temperature before testing. Constructs were submitted for 4-point bending in dorsoplantar, mediolateral, and lateromedial directions. The distance between the external supports of the testing system was 28 cm, and the distance between the internal supports of the testing system was 10 cm. Linear variable differential transformers were placed at the level of the inner supports (Figure 4). A servohydraulic materials-testing machinee was used to evaluate the constructs. For the 4-point bending process, a sinusoidal pattern involving a 10-kN load cell was implemented, which took approximately 4 seconds to load or unload from a 100-N preload to 10 kN, permitting a 5-second pause at 100 N and 10 kN. Data on load (N) and stroke-displacement of the load-ram signals were collected at 10 Hz for data analysis.
Sketch of the placement of an equine tarsus specimen for evaluation in a biomechanical testing machine (A) and photograph of the construct within the testing machine (B). A—The dashed boxes denote the size and shape of the thermal-setting plastic compound. The small dashed circles denote the critical placement of steel alloy rods embedded in the bone and plastic compound. The large arrowheads represent the limits of the outer support, and the small arrowhead represents the application of force to each putty block and the location of the LVDTs.
Citation: American Journal of Veterinary Research 77, 10; 10.2460/ajvr.77.10.1071
With the longitudinal axis of the tarsus aligned to the axis of rotation, torsional tests were performed to analyze for both internal and external rotation by use of a 125-Nm load cell. Constructs were preloaded with 500 N of axial compression, and then torque was applied by use of a sinusoidal pattern, with a 5-second pause at maximum torque and a 10-second pause when the constructs returned to neutral position. In nondestructive axial compression, compression was directed along the longitudinal axis of the tarsus. Each tarsus was loaded under load control by use of a sinusoidal pattern from 500-N preload to 25 kN of maximal compression. A 5-second pause at full loading and a 30-second pause upon unloading to 500 N were used.
For each testing condition, 5 cycles were performed; data for the first 2 preconditioning cycles were discarded and data for the subsequent 3 cycles were used. For the 4-point bending process, bending moments were calculated as load force multiplied by the distance between the outer supports of the bending fixture, divided by 4. Angular displacement (θ) was calculated by use of the following formulae:
where Θ1 and θ2 are the angular rotations at each support, D1 and D2 are the linear displacements of each support by use of the LVDTs, and × is the distance (m) of the LVDT from the pivot point of the inner support (Figure 4).
Bending moment–angular deformation curves were generated for each construct to determine the maximal bending moment and stiffness. Stiffness was determined as the slope of the elastic phase of these curves. For axial compression, construct stiffness was determined from the linear portion of the load-displacement curve. After construct testing, radiographs of each construct were obtained to evaluate for evidence of construct failure.
Statistical analysis
Summary data are reported as mean ± SEM. Data were analyzed for normality of distribution with the Shapiro-Wilk test. Prior to statistical testing, data for test constructs were also normalized to data for the contralateral control limb and reported as a percentage. When data were normally distributed, 1-way ANOVA with repeated measures was performed followed by the Tukey method of multiple comparisons to detect differences among construct groups. When data were nonnormally distributed, the Friedman test was used to detect differences among groups. Values of P < 0.05 were considered significant. Statistical analysis was performed by use of statistical software.f,g
Results
Specimens
Paired tarsi used in the study originated from 14 Quarter Horses, 8 Thoroughbreds, and 2 Standardbreds. Mean ± SEM age of the horses had been 8.9 ± 3.8 years, and mean body weight had been 489.5 ± 42.1 kg. Horses included 2 stallions, 10 geldings, and 12 mares. For tibial blocking, 14 horses had a tibial width < 9.0 cm and 10 had a tibial width ≥ 9.0 cm as measured on dorsoplantar radiographic views.
No significant deformation was identified in any construct or its contralateral control tarsus during testing. Radiographs obtained after biomechanical testing revealed no radiographic evidence of construct failure through either deformation of the implant materials or damage to the bone. Data resulting from each testing modality were normally distributed, so a parametric series of statistical tests was performed.
Biomechanical comparisons normalized to contralateral control tarsi
Across the combination of all testing methods, DKC constructs had significantly (P < 0.01 for all comparisons) greater tarsal stiffness than did their contralateral control tarsi by a mean ± SEM value of 15.7 ± 1.4%. Across all testing methods, DMLCP constructs had significantly (P < 0.01 for all comparisons) greater tarsal stiffness than did their contralateral control tarsi by 12.8 ± 1.5%. No significant difference was identified in these stiffness enhancements between DKC and DMLCP constructs.
Across all testing methods, MLKC constructs did not differ significantly in tarsal stiffness from their control tarsi (mean difference, 0.1 ± 1.6%). The DKC and DMLCP constructs were significantly stiffer than the MLKC constructs. Across all testing methods, surgical drilling significantly (P < 0.01 for all comparisons) weakened the tarsi by reducing their stiffness by 8.2 ± 1.2%. The DKC, DMLCP, and MLKC constructs were significantly superior to the surgical drilling constructs in tarsal stiffness.
Dorsoplantar 4-point bending
For dorsoplantar bending, a significant (P < 0.001) difference was detected among the construct groups (Figure 5). No significant (P = 0.40) effect was identified for the interaction between construct group and cycle, nor was there a significant (P = 0.37) effect of blocking. There were no significant differences in bending stiffness among the contralateral control tarsi for each of the 4 construct groups with respect to dorsoplantar bending. No significant difference (P = 0.58) in bending stiffness was identified between surgical drilling constructs (78,376 ± 462 Nm/radian) and their contralateral control tarsi (83,605 ± 656 Nm/radian). Dorsoplantar bending stiffness of the MLKC constructs (90,524 ± 1,603 Nm/radian) was not significantly (P = 0.06) greater than that of their contralateral control tarsi (85,077 ± 1,227 Nm/radian). The DKC constructs had a significantly (P = 0.002) greater dorsoplantar bending stiffness (106,521 ± 1,281 Nm/radian) than did their contralateral control tarsi (95,439 ± 1,753 Nm/radian). The DMLCP constructs also had a significantly (P < 0.001) greater dorsoplantar bending stiffness (98,235 ± 2,088 Nm/radian) than did their contralateral control tarsi (82,800 ± 1,532 Nm/radian).
Typical load deformation in a representative cycle from progressive nondestructive bending in the dorsoplantar direction, where load is expressed as the bending moment, for DKC (solid black line), DMLCP (solid gray line), MLKC (dotted black line), and surgical drilling (dotted gray line) constructs created in equine tarsus specimens (n = 6/construct type). The slope of the straight region of these curves represents the stiffness. The gradient of the surgical drilling group is less than the gradients of the 3 implant groups (DKC, DMLCP, and MLKC).
Citation: American Journal of Veterinary Research 77, 10; 10.2460/ajvr.77.10.1071
In comparisons among the construct groups, surgical drilling was the weakest in dorsoplantar bending stiffness, representing a 6.2 ± 0.8% reduction relative to the bending stiffness of contralateral control tarsi (Table 1). The surgical drilling constructs were significantly (P < 0.001) weaker in dorsoplantar bending than were the MLKC constructs (4.5 ± 1.6% increase relative to control tarsi), DKC constructs (14.1 ± 1.4% increase), or DMLCP constructs (17.9 ± 3.2% increase). The MLKC constructs were significantly less stiff in dorsoplantar bending than were the DMLCP (P < 0.01) or DKC (P < 0.001) constructs. No significant difference was identified between the DMLCP and DKC constructs in dorsoplantar bending.
Mean ± SEM bending and axial compression values and percentage difference from contralateral control values for surgical drilling, MLKC, DMKC, and DMLCP constructs created from equine tarsus specimens (n = 6/construct or contralateral control group).
| Dorsoplantar bending | Lateromedial bending | Mediolateral bending | Axial compression | |||||
|---|---|---|---|---|---|---|---|---|
| Construct | Value (Nm/radian) | Percentage difference | Value (Nm/radian) | Percentage difference | Value (Nm/radian) | Percentage difference | Value (Nm/radian) | Percentage difference |
| Surgical drilling | 78,376 ± 462 | −6.2 ± 0.8a | 68,974 ± 4,810 | −3.6 ± 1.5a | 80,028 ± 4,382 | −3.6 ± 1.5a | 80,028 ± 4,382 | −12.0 ± 5.0a |
| MLKC | 90,524 ± 1,603 | 4.5 ± 1.6b | 75,033 ± l,832 | 5.6 ± 1.0b | 101,886 ± 1,173 | 5.6 ± 1.0b | 101,886 ± 1,173 | 7.3 ± 3.2a,b |
| DMKC | 106,521 ± 1,281 | 14.1 ± 1.4c | 94,234 ± 2,399 | 14.3 ± 1.4c | 104,371 ± 2,968 | 14.3 ± 1.4c | 104,371 ± 2,968 | 21.2 ± 2.2c |
| DMLCP | 98,235 ± 2,088 | 17.9 ± 3.2c | 78,638 ± 2,962 | 5.4 ± 2.0b,d | 106,769 ± 1,108 | 5.4 ± 2.0b,d | 106,769 ± 1,108 | 21.3 ± 5.0b,c |
Values with different superscript letters within the same column differ significantly (P < 0.05) from each other.
At no point did mean values for surgical drilling constructs surpass those of their contralateral control tarsi (all negative values).
Lateromedial 4-point bending
For lateromedial bending, a significant (P = 0.01) difference was identified among all construct groups. There was no significant (P = 0.16) effect of the interaction between construct group and cycle and no significant (P = 0.77) effect of blocking. No significant differences were identified in lateromedial bending stiffness among the contralateral control tarsi for each of the 4 construct groups. No significant (P = 0.21) difference was identified in bending stiffness between surgical drilling constructs (70,222 ± 1,033 Nm/radian) and their contralateral control tarsi (78,651 ± 1,408 Nm/radian). No significant difference was detected between the MLKC constructs (80,308 ± 1,218 Nm/radian) and their contralateral control tarsi (77,502 ± 1,310 Nm/radian; P = 0.28), between DKC constructs (99,919 ± 1,850 Nm/radian) and their contralateral control tarsi (95,122 ± 3,496 Nm/radian; P = 0.58), or between DMLCP constructs (81,750 ± 2,570 Nm/radian) and their contralateral control tarsi (78,334 ± 2,396 Nm/radian; P = 0.54).
In comparisons among construct groups, surgical drilling constructs were the weakest in lateromedial bending stiffness, representing a 12.4 ± 0.8% reduction relative to the bending stiffness of contralateral control tarsi (Table 1). The surgical drilling constructs were significantly (P < 0.001) weaker in lateromedial bending than were the MLKC constructs (3.8 ± 1.1% increase relative to control tarsi), DKC constructs (11.5 ± 2.3% increase), or DMLCP constructs (4.7 ± 1.8% increase). The MLKC constructs were significantly (P < 0.05) less stiff in lateromedial bending than were the DKC constructs, but not the DMLCP constructs. No significant difference was identified between the DMLCP and DKC constructs in lateromedial bending.
Mediolateral 4-point bending
For mediolateral bending, a significant (P < 0.001) difference was identified among all construct groups. No significant (P = 0.51) effect of the interaction between construct group and cycle was evident, nor was there a significant (P = 0.96) effect of blocking. No significant differences were identified in mediolateral bending stiffness among contralateral control tarsi for each of the 4 construct groups. No significant difference in bending stiffness was detected between surgical drilling constructs (68,974 ± 4,810 Nm/radian) and their contralateral control tarsi (71,839 ± 1,034 Nm/radian; P = 0.99) or between MLKC constructs (75,033 ± 1,832 Nm/radian) and their contralateral control tarsi (75,358 ± 1,442 Nm/radian; P = 1.00). A significant (P = 0.002) difference was identified between DKC constructs (94,234 ± 2,399 Nm/radian) and their contralateral control tarsi (82,841 ± 2,603 Nm/radian). There was no significant (P = 0.16) difference in mediolateral bending stiffness between DMLCP constructs (78,638 ± 2,962 Nm/radian) and their contralateral control tarsi (74,202 ± 1,592 Nm/radian). Among the test specimens within in each group, no significant differences were identified.
In comparisons among construct groups, surgical drilling was the weakest in mediolateral bending stiffness, representing a 3.6 ± 1.5% reduction relative to the bending stiffness of contralateral control tarsi (P < 0.01; Table 1). The surgical drilling constructs were significantly (P < 0.001) weaker in mediolateral bending than were the MLKC constructs (5.6 ± 2.0% increase relative to control tarsi), DKC constructs (14.3 ± 1.4% increase), or DMLCP constructs (5.4 ± 2.0% increase). The MLKC constructs were significantly (P < 0.01) less stiff in mediolateral bending than were the DKC constructs, but not the DMLCP constructs. The DKC constructs were significantly (P < 0.01) stiffer than the DMLCP constructs in mediolateral bending.
In comparisons of the 4-point bending directions, significantly greater stiffness was identified in dorsoplantar bending directions than in lateromedial and mediolateral directions. In addition, constructs bent in the lateromedial direction were significantly (P < 0.01) stiffer than were constructs bent in the mediolateral direction, which was the weakest direction.
Rotational stiffness
For both external and internal rotational stiffness, significant (P ≤ 0.01 for all comparisons) differences were identified among the construct groups. No significant effect of the interaction between construct group and cycle (P > 0.59) or blocking (P > 0.59) was identified in any comparison.
No significant differences in external or internal rotational stiffness were identified among groups of contralateral control tarsi. For the surgical drilling constructs, no significant differences were identified in external (1,156 ± 16 Nm/radian) and internal (1,086 ± 25 Nm/radian) rotation, compared with values for contralateral control tarsi (1,233 ± 21 Nm/radian [P = 0.62] and 1,235 ± 20 Nm/radian [P = 0.13], respectively). For the MLKC constructs, no significant differences were identified in external (841 ± 45 Nm/radian) and internal (996 ± 42 Nm/radian) rotational stiffness, compared with values for contralateral control tarsi (989 ± 23 Nm/radian [P = 0.28] and 1,106 ± 31 Nm/radian [P = 0.92], respectively).
For the DKC constructs, a significantly (P = 0.003) greater external rotational stiffness (1,345 ± 20 Nm/radian) was detected, compared with that for contralateral control tarsi (1,123 ± 34 Nm/radian). No significant (P = 0.90) difference in internal rotational stiffness was identified between DKC constructs (1,257 ± 33 Nm/radian) and their contralateral control tarsi (1,180 ± 65 Nm/radian). For the DMLCP constructs, a significantly (P = 0.002) greater internal rotational stiffness (1,324 ± 20 Nm/radian) was identified, compared with that for contralateral control tarsi (1,152 ± 40 Nm/radian). No significant (P = 0.36) difference in external rotational stiffness was evident between DMLCP constructs (1,263 ± 29 Nm/radian) and their contralateral control tarsi (1,149 ± 28 Nm/radian).
The surgical drilling constructs were weakest in external rotational stiffness (11.2 ± 3.1% reduction in stiffness relative to the rotational stiffness of contralateral control tarsi). The surgical drilling constructs were significantly less stiff in external rotation than were the DKC constructs (11.7 ± 6.0% increase relative to control tarsi [P < 0.01]) or DMLCP constructs (16.6 ± 3.0% increase [P < 0.001]; Table 2). The surgical drilling constructs were not significantly different from the MLKC constructs (8.0 ± 4.8% reduction relative to control tarsi). The MLKC constructs were significantly less stiff in external rotation than were the DMLCP constructs but were not significantly different from the DKC constructs. No significant differences in external rotational stiffness were identified between the DKC and DMLCP constructs.
Mean ± SEM external and internal rotation values and percentage difference from contralateral control values for surgical drilling, MLKC, DMKC, and DMLCP constructs created from equine tarsus specimens (n = 6/construct or contralateral control group).
| External rotation (Nm/radian) | External rotation (°) | Internal rotation (Nm/radian) | Internal rotation (°) | |||||
|---|---|---|---|---|---|---|---|---|
| Construct | Value | Percentage difference | Value | Percentage difference | Value | Percentage difference | Value | Percentage difference |
| Surgical drilling | 1,156 ± 16 | −11.2 ± 3.1a | 9.50 ± 0.45 | 15.8 ± 3.5a | 1,086 ± 25 | −5.9 ± 1.7a | 10.30 ± 0.31 | 9.6 ± 2.5a |
| MLKC | 841 ± 45 | −8.0 ± 3.4a,b | 9.49 ± 0.33 | 19.3 ± 8.5a | 996 ± 42 | 12.9 ± 5.5a,b | 7.96 ± 0.26 | 33.8 ± 10.9a |
| DMKC | 1,345 ± 20 | 11.7 ± 6.0b,c | 7.67 ± 0.33 | −15.9 ± 6.8b | 1,123 ± 34 | 21.5 ± 3.7c | 8.16 ± 0.39 | −20.9 ± 5.1b |
| DMLCP | 1,324 ± 20 | 16.6 ± 3.0c | 9.69 ± 0.47 | −20.6 ± 3.2b | 1,152 ± 40 | 11.2 ± 4.0b,c | 9.68 ± 0.57 | −0.4 ± 4.5a |
See Table 1 for key.
The MLKC constructs were weakest in internal rotational stiffness (12.9 ± 5.5% reduction in stiffness relative to the rotational stiffness of contralateral control tarsi). The MLKC constructs were also significantly (P < 0.001) less stiff in internal rotation than were the DKC constructs (21.5 ± 3.7% increase relative to control tarsi) but did not differ significantly from the DMLCP constructs (11.2 ± 4.0% increase relative to control tarsi) or the surgical drilling constructs (5.9 ± 1.7% decrease relative to control tarsi). The surgical drilling constructs were significantly less stiff in internal rotation than were the DKC constructs but did not differ significantly from the DMLCP constructs. No significant differences in internal rotational stiffness were identified between the DKC and DMLCP constructs.
In terms of rotational stiffness direction, constructs had significantly (P = 0.01) increased stiffness in internal versus external rotation.
Rotational angle
For both external and internal rotational maximum angles as measured at 120 Nm of torque, significant (P ≤ 0.001 for all comparisons) differences were identified among the construct groups. For both the internal and external directions, no significant (P > 0.10) effect of block randomization was identified.
Among the external and internal evaluations, no significant differences were identified among the groups of contralateral control tarsi. For the surgical drilling constructs, no significant differences were identified in external (9.5 ± 0.45°) and internal (10.3 ± 0.31°) rotation, compared with values for contralateral control tarsi (8.2 ± 0.30° [P = 0.93] and 9.4 ± 0.40° [P = 0.46], respectively). The surgical drilling constructs were able to rotate 15.8 ± 3.5% more than their contralateral control tarsi in an external direction and 9.6 ± 2.5% more in an internal direction (Table 2).
For the MLKC constructs, no significant (P = 0.98) difference in external rotation (9.49 ± 0.33°) was identified, compared with that for their contralateral control tarsi (7.96 ± 0.26°). However, for internal rotation, MLKC constructs had a significantly (P = 0.002) greater range of motion (11.99 ± 0.93°) than did control tarsi (8.97 ± 0.11°). The MLKC constructs were able to rotate 19.3 ± 8.5% more than were their contralateral control tarsi in an external direction and 33.8 ± 10.9% more in an internal direction; however, only the increase for internal rotation was significant.
For the DKC constructs, external rotation (7.67 ± 0.33°) did not have significantly (P = 0.69) less freedom than it did for contralateral control tarsi (9.12 ± 0.55°). For internal rotation, DKC constructs had a significantly (P = 0.003) decreased range of motion (8.16 ± 0.39°) relative to that of their contralateral control tarsi (10.32 ± 0.74°). The DKC constructs were able to rotate 15.9 ± 6.8% less than were control tarsi in an external direction (P = 0.09) and 20.9 ± 5.1% less in an internal direction (P = 0.008).
For the DMLCP constructs, a significant (P < 0.001) difference was identified for external rotation (9.69 ± 0.47°), compared with values for contralateral control tarsi (12.2 ± 1.05°). For internal rotation, no significant (P = 0.14) difference was identified between DMLCP constructs (9.68 ± 0.57°) and their contralateral control tarsi (9.64 ± 0.32°). The DMLCP constructs were able to rotate 20.6 ± 3.2% less than were control tarsi in an external direction and 0.4 ± 4.5% more in an internal direction; however, only the difference for the external direction was significant.
The surgical drilling and MLKC constructs were able to rotate significantly (P < 0.05 for all comparisons) more in an external direction than were the DKC or DMLCP constructs. No significant difference in degrees of external rotation was identified between the surgical drilling and MLKC constructs, nor was any such difference identified between the DKC and DMLCP constructs. There was no significant difference among surgical drilling, MLKC, and DMLCP constructs in degrees of internal rotation. The DKC constructs had significantly fewer degrees of rotation in an internal direction than did the MLKC, surgical drilling, and DMLCP constructs.
Axial compression
For axial compression, a significant (P = 0.003) difference was identified among all construct groups (Table 1). No significant (P = 0.08) effect was identified for the interaction between construct group and cycle; however, a significant (P = 0.04) blocking effect was identified, indicating that horses with tibial mediolateral diameters ≥ 9.0 cm as measured at the physeal scar had significantly greater tarsal axial stiffness than did horses with tibial measurements < 9.0 cm. There were no significant differences in axial stiffness among the 4 groups of contralateral control tarsi.
No significant (P = 0.09) difference in axial compression was identified between the surgical drilling constructs (80,028 ± 4,382 N/mm) and their contralateral control tarsi (92,054 ± 1,173 N/mm), nor was there a significant (P = 0.97) difference between the MLKC constructs (101,886 ± 1,806 N/mm) and their contralateral control tarsi (96,600 ± 3,628 N/mm). The DKC test constructs had significantly (P = 0.003) greater axial stiffness (104,371 ± 2,968 N/mm) than did their contralateral control tarsi (86,338 ± 2,409 N/mm). The DMLCP test constructs had significantly (P = 0.03) greater axial stiffness (106,769 ± 1,108 N/mm) than did their contralateral control tarsi (90,375 ± 3,277 N/mm).
In comparisons among the construct groups, surgical drilling was the weakest in axial compression, with a 12.0 ± 5.1% reduction in compression stiffness relative to that of contralateral control tarsi. The surgical drilling constructs were significantly (P < 0.001) weaker in axial compression than were the DKC constructs (21.2 ± 2.2% increase relative to control tarsi) or DMLCP constructs (21.3 ± 5.0% increase). There was no significant difference in axial compression stiffness between the surgical drilling and MLKC constructs, nor was a difference identified between the MLKC and DMLCP constructs, although the DKC constructs were significantly stiffer than the MLKC constructs. No significant difference was identified between DMLCP and DKC constructs in axial compression stiffness.
Discussion
In the study reported here, biomechanical properties of various fixation methods for the distal aspect of the tarsus were evaluated. To the authors’ knowledge, this represents the first study in which the biomechanical properties of these methods were assessed. Previous studies have focused on either arthrodesis of joints21–24 with a higher degree of motion than tarsal joints have or on the biomechanical properties of diaphyseal sections of equine bones.25
In terms of bending stiffness, arthrodeses of high-motion joints are reported to have bending stiffness values of approximately 1,000 Nm/radian.26,27 In a study25 involving diaphyseal bones, investigators calculated that the stiffness of MT3 was approximately 13,000 Nm/radian for 4-point bending, 2,000 Nm/radian for torsion of the femur, and 6,000 N/mm for axial compression. The present study revealed that the bending stiffness in the centrodistal and tarsometatarsal joints was approximately 90,000 Nm/radian for dorsoplantar bending, 1,000 Nm/radian for rotational stiffness, and 90,000 N/mm for axial compression.
The values derived in the present study can potentially be explained if one considers the area moment of inertia (I), which relates resistance to bending, and the polar moment of inertia (J), which relates resistance to torsional deformation. For a hollow tube, the following relationship exists25:
where Ix is the area moment along the x-axis, r2 is the radius of the outer diameter, and r1 is the radius of the inner diameter. Area moments of inertia are often identified along the x-, y-, and z-axes, depending on the area of interest, and are dependent on mass, density, and volume. Polar moments of inertia are given by the following formula25:
where D is the outer diameter and d is the inner diameter. Therefore, the sum of second moments of area about orthogonal axes is a function only of the position of the origin O for the axes. The polar moment of area about the origin (IO) is calculated as follows25:
where Ixx and Iyy are the polar moments of inertia in the x- and y-axes, respectively.
From these equations, it can be understood that resistance to bending or torsion is related to the fourth power of the diameter. Therefore, for a 2-times increase in diameter, there is a 16-times increase in the moment of inertia. Not only are the diameters of the centrodistal and tarsometatarsal joints larger than that of MT3, the involved bones have a solid core and not a hollow cylinder like the diaphysis of MT3. The cortical density of MT3 may also be greater than that of the bones in the distal aspect of the tarsus; therefore, this must also be taken into account. However, in the 4-point bending scheme, if the centrodistal and tarsometatarsal joints are considered as being of similar density to and hollow like MT3, but with twice the diameter, the bending moment could be expected to be 16 times that obtained in the other study,25 placing the bending moment at approximately 208,000 Nm/radian. If this value is used as a baseline, taking into account the reduced bone density of the centrodistal and tarsometatarsal joints and the nonhollow properties, combined with the testing of 3 joint levels, then the value of approximately 90,000 Nm/radian obtained in the present study appears rational.
In horses with clinical osteoarthritis, the appearance of osteophytes occurs mainly on the dorsal aspect of the joints and is most notable on lateromedial radiographs.28 The preferential deposition of bone on the dorsal surface may be a functional adaptation toward neutralizing the forces exerted across the dorsal aspect of the tarsus during loading. The common finding of osteophytes in the dorsal aspect of the centrodistal and tarsometatarsal joints is clinical evidence that supports the relevance of the biomechanical findings that the greatest bending moment of the specimens used in the present study was in the dorsoplantar direction, suggesting that this region is the most important in maintaining stability across the centrodistal and tarsometatarsal joints. Alternatively, the preference for osteophyte formation on the dorsal aspect may be due to the extensive ligamentous insertions in this area, predisposing the dorsal aspect to enthesiophyte formation. A combination of both these potential contributing factors is likely to be the most plausible reason for osteophyte formation in this location of the tarsus. Growth of osteophytes in this region may also serve to increase the cross-sectional area of the bone, resulting in a modification of the moment of inertia for the bone and thus affecting the stiffness in a bending direction.
The same principles would apply to rotational torsion, whereby, given the same parameters, we would expect the rotational stiffness to increase 16-fold relative to that for MT3, from approximately 800 Nm/radian to 12,800 Nm/radian. However, values of approximately 1,000 Nm/radian were obtained in the present study. Although this value is conveniently between the rotational stiffness of the tibia (approx 1,200 Nm/radian) and MT3 (approx 800 Nm/radian), it needs to be explained in light of the calculation that the polar moment of inertia would provide.
In the rotational method of analysis, rotational forces are exerted across the joint. This inevitably causes a combination of both rotation and translation to occur about the centrodistal and tarsometatarsal joints. Therefore, one would expect more movement to occur in this plane than in any of the 4-point bending tests performed. In the other study25 involving equine diaphyseal bones, rotational torque tests were stopped when grip slippage occurred, which was at 1.6° for the femur and 2.2° for MT3. No slippage occurred in any constructs evaluated in the present study, which were rotated approximately 10°. Therefore, considerably more laxity existed in the centrodistal and tarsometatarsal joints than in the diaphysis of the MT3. Given these findings, our value of 1,000 Nm/radian is within a reasonably expected value for tarsal rotation.
For axial compression, the surface area of compression for diaphysis of long bones must be considered in comparison with that of the distal aspect of the tarsus. In the diaphysis, diameter ratios for cortical bone in other species can be approximately 25%.25 Thus, the surface area for the joints used in the present study, which was twice the diameter of MT3 and nonhollow, was approximately 64-fold as great as that of the cortex at the diaphysis of MT3. If we therefore normalized the values obtained for compressive stiffness to the area of bone, the results would be within the range of the axial compression values previously attained for MT3. A significant effect of tibial width was identified in the present study, with horses with larger tibiae having significantly greater axial stiffness than horses with smaller tibiae. In general, the effect of contact area and congruency across the joint surfaces may result in some amount of destabilization.
In relation to the hypotheses tested in the present study, surgical drilling resulted in a reduction in tarsal stiffness by approximately 8.2% across all tests performed. This could potentially have important consequences when considering the effects of this surgical technique on the recovery period for horses after surgery. Biomechanical testing also revealed that equine centrodistal and tarsometatarsal joints had the greatest resistance to bending in the dorsoplantar direction than in other directions and that they were stiffer in the lateromedial direction than in the mediolateral direction. Because the orientation of the trochlear ridges is moderately lateral, resulting in abduction of the distal portion of the limb during ambulation and flexion, the medial aspect of the tarsus could be expected to be the weakest portion of the tarsus when evaluated biomechanically. The medial collateral ligaments are reportedly more prone to injury in horses than the lateral collateral ligaments,29 and the data reported here provide a possible explanation for this clinical phenomenon.
Consistent with our second hypothesis, arthrodesis by use of DKC or DMLCP methods significantly increased the stiffness of the centrodistal and tarsometatarsal joints in 4-point bending, rotation, rotational angle limit, and axial compression, compared with values for contralateral control tarsi. Across all the methods of testing, DKC constructs had greater tarsal bending stiffness by approximately 15.7% and DMLCP constructs by approximately 12.8%. However, we were unable to identify any significant differences between the DKC and DMLCP constructs. Given that the DKC construct was similar in stiffness to the DMLCP construct, yet can be placed in a less invasive manner, clinical evaluation of this technique may be warranted.
Results of the study reported here indicated that dorsal placement of a single kerf-cut cylinder was superior to mediolateral placement of 2 cylinders. Stainless steel cylinders were used in a previous study19 for tarsal arthrodesis, whereby 2 cylinders were placed at each joint level.19 Although such use was successful, 1 horse in that study subsequently sustained a fracture of the third tarsal bone. This was believed to have occurred as a result of incorrect placement of the implants, emphasizing the importance of the surgical technique. The present study involved a similar method of placement for MLKC constructs; however, a kerf-cut cylinder was used that is more rigid in fixation to the bone (Figure 1). In general over all biomechanical testing methods, MLKC placement resulted in a 0.1% difference from control values. Therefore, we were unable to conclude that the use of MLKC provided a significant improvement in overall tarsal stability. Interestingly, although it appeared that placement of the MLKC enhanced tarsal stability during the 4-point bending tests, MLKC constructs were weaker than their contralateral control tarsi when evaluated in rotational analysis. Placement of the DKC, however, resulted in a significant increase in tarsal stiffness over all biomechanical testing methods. Therefore, placement of a single, larger DKC that spanned both the centrodistal and tarsometatarsal joint spaces was biomechanically superior.
In the present study, axial compression of 25,000 N was used, which is a force 5 times the body weight of a typical horse, with no adverse effects and no fracturing during testing of equine tarsus specimens. However, a limitation of this study was that the method of loading used was slower and more gradual rather than the explosive forces generated when horses exercise or trot. Furthermore, tarsus specimens without overt evidence of osteoarthritis were used to minimize variation. It is highly probable that tarsi of horses with clinical evidence of osteoarthritis would have different biomechanical properties because of joint space narrowing and the presence of bridging osteophytes. Therefore, caution should be exercised when considering the biomechanical results reported here in a clinical context.
To further stabilize joints during arthrodesis, recommendations are to provide compression across joint surfaces. An additional limitation of the study reported here was that the DMLCP constructs involved no placement of transarticular screws. The aim was to span the joints and provide a locking construct to ensure security. It is entirely conceivable that the construct would have been stronger if a series of screws had been placed across the centrodistal and tarsometatarsal joints to apply a compressive force across the joint surfaces. Whether the forces exerted across the joints during the rotational evaluation would have exerted enough stress to bend or break the transarticular screws remains unknown. Because the order of tests was randomized, we believed it wise to avoid damaging the constructs during the rotational testing.
Results of the present study suggested that surgical drilling of equine centrodistal and tarsometatarsal joints may destabilize the biomechanical properties of that region, although those findings were not significant. Use of a single DKC or DMLCP to enhance arthrodesis of the centrodistal and tarsometatarsal joints resulted in an immediate improvement in stiffness. The ability to place a DKC without extensive soft tissue dissection and disruption in the area therefore suggested that clinical application of this particular implant would be worthy of investigation.
Acknowledgments
Supported by a grant from the American College of Veterinary Surgeons Surgeon-in-Training program and the Companion Animal Fund Program, School of Veterinary Medicine, University of Wisconsin-Madison.
Presented in abstract form at the American College of Veterinary Surgeons Symposium, San Antonio, Tex, 2013.
The authors thank Ryan Buck and Synthes for assistance with and provision of the DMLCPs; Ron McCabe and Ray Vanderby Jr. for biomechanical assistance; and Yan Lu, Carissa Sawyer, Sarah Rossmiller, and Alyssa White for other assistance.
ABBREVIATIONS
| DKC | Dorsally applied kerf-cut cylinder |
| DMLCP | Dorsomedially applied locking compression plate |
| LVDT | Linear variable differential transformer |
| MLKC | Medially to laterally placed kerf-cut cylinder |
| MT3 | Third metatarsal bone |
Footnotes
Department of Engineering, University of Wisconsin-Madison, Madison, Wis.
Wilson Tool and Manufacturing, Spokane, Wash.
Ti/3H 4.5–5.0 mm, model No. 440.131, Synthes, Zuchwil, Switzerland.
Bondo, 3M, Saint Paul, Minn.
Instron 1350 universal testing machine, Instron Corp, Canton, Mass.
SAS, version 9.4, SAS Institute Inc, Cary, NC.
Prism, version 6, GraphPad Software Inc, La Jolla, Calif.
References
1. McIlwraith C, Robertson J. Arthrodesis of the distal tarsal joints. Equine Surg Adv Techniques 1987:185–190.
2. Kidd J, Fuller C, Barr A. Osteoarthritis in the horse. Equine Vet Educ 2001;13: 160–168.
3. Rice JR, Pisetsky DS. Pain in the rheumatic diseases: practical aspects of diagnosis and treatment. Rheum Dis Clin North Am 1999;25: 15–30.
4. Heijink A, Gomoll AH, Madry H, et al. Biomechanical considerations in the pathogenesis of osteoarthritis of the knee. Knee Surg Sport Traumatol Arthrosc 2012;20: 423–435.
5. Simmons E, Bertone A, Weisbrode S. Instability-induced osteoarthritis in the metacarpophalangeal joint of horses. Am J Vet Res 1999;60: 7–13.
6. Rooney JR, Turner LW. The mechanics of horses pulling loads. J Equine Vet Sci 1985;5: 355–359.
7. Baxter GM, Dechant JE, Southwood LL. Diagnosis of distal tarsal osteoarthritis in horses. Compend Contin Educ Vet 2003;25: 138–147.
8. Edwards G. Surgical arthrodesis for the treatment of bone spavin in 20 horses. Equine Vet J 1982;14: 117–121.
9. Moyer W, Brokken T, Raker C. Bone spavin in Thoroughbred race horses, in Proceedings. Ann Conv Am Assoc Equine Pract (USA) 1983.
10. Dechant JE, Baxter GM, Southwood LL, et al. Use of a three-drill-tract technique for arthrodesis of the distal tarsal joints in horses with distal tarsal osteoarthritis: 54 cases (1990–1999). J Am Vet Med Assoc 2003;223: 1800–1805.
11. Barber SM. Arthrodesis of the distal intertarsal and tarsometatarsal joints in the horse. Vet Surg 1984;13: 227–235.
12. Zubrod CJ, Schneider RK, Hague BA, et al. Comparison of three methods for arthrodesis of the distal intertarsal and tarsometatarsal joints in horses. Vet Surg 2005;34: 372–382.
13. Shoemaker RW, Allen AL, Richardson CE, et al. Use of intra-articular administration of ethyl alcohol for arthrodesis of the tarsometatarsal joint in healthy horses. Am J Vet Res 2006;67: 850–857.
14. Wyn-Jones G, May S. Surgical arthrodesis for the treatment of osteoarthrosis of the proximal intertarsal, distal intertarsal and tarsometatarsal joints in 30 horses: a comparison of four different techniques. Equine Vet J 1986;18: 59–64.
15. Scruton C, Baxter G, Cross M, et al. Comparison of intra-articular drilling and diode laser treatment for arthrodesis of the distal tarsal joints in normal horses. Equine Vet J 2005;37: 81–86.
16. von Salis B, Auer JA, Fackelman GE. Small tarsal joint arthrodesis. In: Fackelman GE, Auer JA, Nunamaker DM, eds. AO principles of equine osteosynthesis. New York: Thieme, 2000; 269–279.
17. Wagner PC, Grant BD, Bagby GW, et al. Evaluation of cervical spinal fusion as a treatment in the equine “wobbler” syndrome. Vet Surg 1979;8: 84–88.
18. Crawley GR, Grant BD, White KK, et al. A modified Cloward's technique for arthrodesis of the normal metacarpophalangeal joint in the horse. Vet Surg 1988;17: 117–127.
19. Archer R, Schneider R, Lindsay W, et al. Arthrodesis of the equine distal tarsal joints by perforated stainless steel cylinders. Equine Vet J 1988;20: 125–130.
20. Reardon R, Bailey R, Walmsley J, et al. A pilot in vitro biomechanical comparison of locking compression plate fixation and kerf-cut cylinder fixation for ventral fusion of fourth and fifth equine cervical vertebrae. Vet Comp Orthop Traumatol 2009;22: 371–379.
21. Sod GA, Mitchell CF, Hubert JD, et al. In vitro biomechanical comparison of equine proximal interphalangeal joint arthrodesis techniques: prototype equine spoon plate versus axially positioned dynamic compression plate and two abaxial transarticular cortical screws inserted in lag fashion. Vet Surg 2007;36: 792–799.
22. Knox P, Watkins J. Proximal interphalangeal joint arthrodesis using a combination plate-screw technique in 53 horses (1994–2003). Equine Vet J 2006;38: 538–542.
23. Read EK, Chandler D, Wilson DG. Arthrodesis of the equine proximal interphalangeal joint: a mechanical comparison of 2 parallel 5.5 mm cortical screws and 3 parallel 5.5 mm cortical screws. Vet Surg 2005;34: 142–147.
24. Wolker RR, Carmalt JL, Wilson DG. Arthrodesis of the equine proximal interphalangeal joint: a biomechanical comparison of two parallel headless, tapered, variable-pitched, titanium compression screws and two parallel 5.5 mm stainless-steel cortical screws. Vet Surg 2009;38: 861–867.
25. Hanson P, Markel M, Vanderby R Jr. Diaphyseal structural properties of equine long bones. Am J Vet Res 1995;56: 233–240.
26. Watt BC, Edwards RB, Markel MD, et al. Arthrodesis of the equine proximal interphalangeal joint: a biomechanical comparison of two 7-hole 3.5-mm broad and two 5-hole 4.5-mm narrow dynamic compression plates. Vet Surg 2002;31: 85–93.
27. Watt BC, Edwards RB, Markel MD, et al. Arthrodesis of the equine proximal interphalangeal joint: a biomechanical comparison of three 4.5-mm and two 5.5-mm cortical screws. Vet Surg 2001;30: 287–294.
28. Fairburn A, Dyson S, Murray R. Clinical significance of osseous spurs on the dorsoproximal aspect of the third metatarsal bone. Equine Vet J 2010;42: 591–599.
29. Sherlock C, Eggleston R, Peroni J, et al. Desmitis of the medial tarsal collateral ligament in 7 horses. Equine Vet Educ 2012;24: 72–80.
