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. Author manuscript; available in PMC: 2018 Aug 16.
Published in final edited form as: J Biomech. 2017 Jul 25;61:160–167. doi: 10.1016/j.jbiomech.2017.07.018

Ultrashort echo time T2* values decrease in tendons with application of static tensile loads

Saeed Jerban 1,*, Amin Nazaran 1, Xin Cheng 1, Michael Carl 2, Nikolaus Szeverenyi 1, Jiang Du 1, Eric Y Chang 3,1
PMCID: PMC5581207  NIHMSID: NIHMS895573  PMID: 28780188

Abstract

In early stages of tendon disease, mechanical properties may become altered prior to changes in morphological anatomy. Ultrashort echo time (UTE) magnetic resonance imaging (MRI) can be used to directly detect signal from tissues with very short T2 values, including unique viscoelastic tissues such as tendons. The purpose of this study was to use UTE sequences to measure T2*, T1 and magnetization transfer ratio (MTR) variations of tendon samples under static tensile loads. Six human peroneal tendons were imaged before and under static loading using UTE sequences on a clinical 3T MRI scanner. Tendons were divided into two static tensile loading groups: group A that underwent one-step loading (15 N) and group B that underwent two-step loading (15 and 30 N). The T2*, T1 and MTR variations were investigated in two selected ROIs, including whole and core sections. Mean T2* values for the first step of loading (groups A and B) in both whole section and core section regions of interest (ROIs) were significantly decreased by 13 ± 7% (P=0.028) and 16 ± 5% (P=0.017), respectively. For the second loading step (group B), there was a consistent, but non-significant reduction in T2* value by 9 ±2% (P=0.059) and 7 ± 5% (P=0.121) for whole and core sections, respectively. Mean T1 did not show any consistent changes for either loading steps (P>0.05). Mean MTR increased slightly, but not significantly for both loading steps (P>0.05). Significant differences were found only in T2* values of tendons by static tensile load application. Therefore, T2* monitoring during loading is suggested for quantitative investigation of the tendons biomechanics.

Keywords: tendon, tensile loading, MRI, ultrashort TE, T2*

1. Introduction

Tendons are unique viscoelastic connective tissues that are responsible for transferring the mechanical loads generated by muscles to the bones (Aparecida de Aro et al., 2012; Franchi et al., 2007b). Water, as the main viscous component (Wellen et al., 2005) of tissues, may comprise over 60% of the total tendon weight (Aparecida de Aro et al., 2012)). Of the dry weight of the tendon, 60–85% is composed of highly organized collagenous fibers, particularly collagen type I (Aparecida de Aro et al., 2012). Collagenous fibers count for the elastic component of tendons (Wellen et al., 2005)). The mechanical properties and great strength of tendons can be partially explained by their highly organized and hierarchical structure, which consist of collagen fibrils, fibers, bundles of fibers and fascicles (Aparecida de Aro et al., 2012; Khan et al., 1999). Tendons demonstrate crimp or zigzag patterns in their microstructure that enable them to withstand, up to 3% strain without damage (Aparecida de Aro et al., 2012; Hansen et al., 2002; Khan et al., 1999).

Magnetic resonance (MR) imaging, as a non-invasive modality, is routinely used for the diagnosis of tendon disease (Pierre-Jerome et al., 2010). The detected signal intensity of a tissue in MR imaging depends on various factors including relaxation time of transverse magnetization (T2) (Chang et al., 2015b; Hendee and Morgan, 1984). Tendons possess a high percentage of organized collagenous matrix that results in very short T2 (Chang et al., 2015b). Therefore, tendon is represented as very low signal intensity when imaged with conventional clinical MR sequences, similar to background noise (Chang et al., 2015b; Du et al., 2010b; Juras et al., 2012). Specifically, the tendon transverse magnetization decays faster than the capability of conventional MRI sequences to capture the signal. However, by using ultrashort echo time (UTE) imaging techniques, signal can be detected in tendons (Chang et al., 2015b; Du et al., 2010b; Jerban et al., 2017; Juras et al., 2013, 2012; Nazaran, 2016; Nazaran et al., 2017). In UTE-MRI, signal can be acquired at very short echo times (on the order of tens of microseconds), before complete decay of transverse magnetization (Chang et al., 2015b; Du et al., 2010b; Juras et al., 2012). In addition to improved visualization, UTE-MRI allows quantitative analysis of the MRI-based properties of tendons, such as apparent relaxation time of transverse magnetization (T2*), relaxation time of longitudinal magnetization (T1), and magnetization transfer (MT) measurements such as MT ratio (MTR) (Chang et al., 2015b, 2014a, 2014b, Du et al., 2010a, 2010b, Juras et al., 2013, 2012; Koff et al., 2014; Syha et al., 2014). Such quantitative UTE biomarkers are hypothesized to reflect histologic changes such as degeneration, including collagen disorganization and increases in total water content (Juras et al., 2013, 2012; Khan et al., 1999; Nazaran et al., 2017; Syha et al., 2014).

Early stage diagnosis of tendon diseases is of great interest to the musculoskeletal research community. Different diseases are hypothesized to alter the mechanical performance of tissues earlier than changes in their anatomical morphology. Hence, researchers have attempted to monitor the changes in MR imaging biomarkers after tensile loading in tendon to assess the sensitivity to mechanical alterations (Han et al., 2000; Helmer et al., 2006, 2004; Koff et al., 2014; Mountain et al., 2011; Shao et al., 2016; Syha et al., 2014; Wellen et al., 2005, 2004), however, only a few of them have used UTE-MRI (Koff et al., 2014; Shao et al., 2016; Syha et al., 2014).

During tendon stretching, two main phenomena occur; first, the number of crimps decrease and the mean crimp angle flattens (Franchi et al., 2007a, 2007b; Legerlotz et al., 2014; Morgan et al., 2006; Mountain et al., 2011), and second, free water redistributes within (Helmer et al., 2004; Wellen et al., 2005, 2004) and outside the tissue (Han et al., 2000; Hannafin and Arnoczky, 1994). In addition, the diameter of larger collagen fibrils decreases and less interspace areas can be found in loaded tendons (Morgan et al., 2006). Therefore, MRI properties of tendons may be expected to change under mechanical loading, even though reported changes in the literature have been inconsistent and are not well understood.

Among the few reported UTE-MRI based studies of tendons under the mechanical loading, Syha et al. observed significant increase of UTE-MTR in vivo (i.e., off-resonance radiofrequency saturation ratio, ORS) for tendons, after short-term exercises (Syha et al., 2014). Chang et al. (Chang et al., 2015a) also observed a significant increase in UTE-MTR in human tendon under static tensile loads (5–10 N). Koff et al. used UTE-MRI for imaging of cyclically loaded (100 cycles of 45 N) rabbit patellar tendon (Koff et al., 2014). The tendons presented significantly shorter T2* values after cyclic loading. In our previous study, (Chang et al., 2015c), UTE-MRI was used for assessing statically loaded human tendons by 10 N loads. No significant changes in UTE biomarkers were observed using single- and bi-component T2* analyses. The applied load in our previous study (Chang et al., 2015c), may have been too low to exert significant changes in the tendons, similar to the fact that Helmer et al. (Helmer et al., 2004) did not observe changes in rabbit Achilles tendons when only 5 N tensile load was applied.

The main objective of this study was to determine whether the application of higher static tensile loads (>10 N) on human tendons alter T2*, T1, and MTR as measured using UTE sequences.

2. Material and methods

2.1. Sample preparation

Six peroneal tendon samples (longus and brevis) were harvested from the ankles of three donors (all female; 38-, 40- and 95-year-old). Tendon samples were cut into 6 cm length and then distributed into two groups. One sample from each donor (longus or brevis) was selected randomly for group A to undergo a one-step loading protocol (15 N). The three remaining samples were assigned to group B and underwent a two-step loading protocol (15 and 30 N). Tendon samples were subjected to two cycles of freezing and thawing. All samples were immersed in saline for 2 hours at 4°C and gently mounted into a plastic loading device fabricated in-house. During the mounting and installing process, the samples were kept hydrated using saline. The loading device chamber was filled with perfluoropolyether (Fomblin, Ausimont, Thorofare, NJ) to minimize dehydration and susceptibility artifact.

2.2. Mechanical loading

An MRI compatible loading device was fabricated using a plastic spring (LL075100U40G, Lee Spring, Brooklyn, NY). The spring was compressed by a manually driven screw, mounted on a notched standard 1×6-inch polyvinyl chloride (PVC) pipe (Figure 1a). Two sides of the tendon specimens were clamped in “fixed” and “sliding” grippers placed on the notched pipe. The two ends of the tendon were fixed firmly in the grippers using two sets of 1/4-inch Nylon bolts and nuts (Figure 1b). The applied tensile mechanical load was calculated based on the amount of spring deformation and its stiffness. The stiffness of springs was calculated by the manufacturer using a load tester machine (UT 11, Larson Systems Inc., Minneapolis, MN, USA) after four loading sessions from 0 to maximum compression (i.e., 2.5, 5.0, 7.5, 10.0, and 15 mm). This was aimed to reduce the initial plastic creep, to achieve an accurate stiffness value and, more importantly, to reduce stiffness changes along the loading experiment in MRI. Along the four loading sessions, the average calculated stiffness converged from 2.10 to 1.98 N/mm.

Figure 1.

Figure 1

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MRI compatible, plastic, loading device used for applying static extension loads on peroneal tendons soaked in fomblin. (b) Mounted tendon on the fixed and sliding grippers, using plastic bolts.

For the scans during the unloaded stage, the screw was driven gently until the beginning of spring deformation (1 N approx.) to make sure that the tendon samples were straight. For the scans at loading stage, the load was adjusted after 2 minutes approximately prior to scanning to compensate the limited relaxation of the tendon.

2.3. UTE-MRI sequences

The loading device and specimens were placed parallel to the B0 field and imaged in the axial plane on a clinical 3T MRI scanner (Signa HDx, GE Healthcare Technologies, Milwaukee, WI) using a wrist coil (BC-10, Medspira, Minneapolis, MN). A quantitative imaging protocol was performed, consisting of I) dual-echo 3D-UTE-Cones sequences (TR = 24.3 ms, TEs = 0.032, 0.4, 0.8, 4.4, 8.8, 13 ms) for T2*measurements, II) variable TR 3D-UTE-Cones sequences (TE = 0.032 ms, TRs = 7.3, 15, 30, 50, 80 ms) for T1 measurements, and III) 3D-UTE-MT-Cones sequences (MT pulse power = 1500°, frequency offset = 2, 5, 10 kHz plus one acquisition without MT preparation) for MTR calculations. Total scan time was 8:36, 20:08, and 7:28 for dual-echo 3D-UTE, variable TR 3D-UTE, and 3D-UTE-MT sequences, respectively. Other imaging parameters include field of view (FOV) = 4 cm, acquisition matrix = 192×192, slice thickness = 3 mm, number of slices = 20, flip angle=10° and receive bandwidth = 62.5 kHz.

2.4. Data analysis

For each dataset, two regions of interest (ROIs) were selected covering the whole and core regions of the tendons (Figure 2). Single-component fitting models were utilized for T2* decay (S(TE) ∝ exp (−TE/T2*) + constant)) and T1 recovery (S(TR) ∝ 1 − exp (−TR/T1) + constant) analyses. The MTR for each off-resonance frequency was calculated by the following equation: MTR = (S0 − SSAT)/S0, where S0 denotes the mean signal intensity acquired without a saturation pulse and SSAT denotes the mean signal intensity acquired with the saturation pulse. All analyses were performed using an in-house developed code in MATLAB (Mathworks, Natick, MA, USA). The T2*, T1 and MTR values as well as their pixel maps in loaded tendon samples were compared with the same samples before loading.

Figure 2.

Figure 2

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UTE-MR images with T2* decay curves comparing the three loading conditions in sample B-1. Single component fitting of T2* decay in the whole and the core sections for (a,b) non-loaded (straight), compared with (c,d) 15 and (e,f) 30 N loaded tendons.

2.5. Statistical Analysis

All statistical analyses were performed using a statistical programming language (R, version 3.2.5, R Development Core Team, Vienna, Austria). The average values of T2*, T1 and MTR in unloaded and loaded tendons were compared using the two-tailed paired t-student test. P-values below 0.05 were considered significant.

3. Results

Figure 2 shows the single component exponential decay fitting at the three loading conditions (non-loaded, 15, and 30 N, respectively) for two selected ROIs (whole section and the core section) of representative sample B-1. The fitting curves demonstrated very good agreements with the actual data for both of the ROIs. The calculated T2* values showed a decrease for both the whole and the core sections after application of the first loading step (15 N), however the T2* reduction was less after the second step of loading (30 N).

The estimated T2* of all examined tendons at different loading steps, based on the single component fitting analyses within the whole and the core sections, are presented in Table 1. For all samples in group A and group B, the T2* values were reduced by applying mechanical loads.

Table 1.

T2* values from single component decay analyses for all the samples (groups A and B) at different loading steps within the whole and the core sections of tendons.

T2* (ms)
Sample
No		 ROI Unloaded Load 1 Load 2
A-1 Whole 0.47±0.03 0.46±0.03 NA
Core 0.51±0.04 0.47±0.03 NA
A-2 Whole 1.04±0.08 0.89±0.07 NA
Core 1.12±0.08 0.90±0.07 NA
A-3 Whole 0.69±0.12 0.68±0.11 NA
Core 0.59±0.09 0.52±0.06 NA
B-1 Whole 1.02±0.13 0.82±0.09 0.77±0.08
Core 1.07±0.08 0.86±0.07 0.82±0.07
B-2 Whole 0.47±0.04 0.38±0.03 0.34±0.02
Core 0.46±0.03 0.37±0.03 0.32±0.03
B-3 Whole 1.23±0.21 1.01±0.19 0.91±0.17
Core 1.53±0.20 1.22±0.18 1.15±0.17
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Table 2 presents the estimated T1 values, based on the single component fitting analyses within the whole and the core sections of the tendons, at different loading steps. No clear trend for T1 value change was observed, by applying mechanical tensile load. However, for five of the scanned samples, T1 values were increased at the first loading step.

Table 2.

T1 values of all the scanned tendon samples (groups A and B) at different loading steps within the whole and the core sections of tendons.

T1 (ms)
Sample No ROI Unloaded Load 1 Load 2
A-1 Whole 610±30 740±40 NA
Core 640±20 760±40 NA
A-2 Whole 840±20 840±30 NA
Core 830±30 840±30 NA
A-3 Whole 590±10 630±10 NA
Core 580±10 590±10 NA
B-1 Whole 900±20 1140±30 810±10
Core 870±30 1110±40 800±20
B-2 Whole 590±60 570±20 580±20
Core 590±80 540±20 590±20
B-3 Whole 590±30 620±50 610±10
Core 660±60 700±70 670±10
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MTR of all the scanned tendon samples (groups A and B) at 2, 5 and 10 kHz off-resonance frequencies for 1500° MT pulse, are presented in Table 3. The MTR change did not show a consistent trend for all scanned tendon samples. However, for most of the cases there was an increase in MTR.

Table 3.

Magnetization transfer ratio (MTR) of all the scanned tendon samples (groups A and B) at different loading steps within the whole and the core sections. The MTR is calculated at 2, 5 and 10 kHz off-resonance frequencies for 1500 deg MT power.

MTR (%)
Unloaded Load 1 Load 2
Sample No ROI 2 kHz 5 kHz 10 kHz 2 kHz 5 kHz 10 kHz 2 kHz 5 kHz 10 kHz
A-1 Whole 16.6 12.0 8.6 17.7 12.3 9.0 NA NA NA
Core 15.8 11.5 8.2 17.8 12.3 9.2 NA NA NA
A-2 Whole 29.6 22.9 17.6 31.3 23.8 18.2 NA NA NA
Core 30.2 23.2 17.5 31.8 24.2 18.3 NA NA NA
A-3 Whole 16.9 11.2 7.3 16.4 12.3 11.0 NA NA NA
Core 22.8 16.1 11.5 22.8 16.6 12.1 NA NA NA
B-1 Whole 22.4 17.6 14.0 23.7 17.6 14.4 26.0 18.7 15.6
Core 24.1 19.3 15.3 24.9 18.1 14.8 26.8 20.0 16.1
B-2 Whole 29.7 20.2 14.9 29.7 19.0 14.6 28.0 18.7 14.2
Core 32.2 21.2 16.5 32.7 21.0 15.8 32.0 21.9 16.1
B-3 Whole 16.8 12.0 9.4 16.9 12.3 9.5 17.2 12.7 9.4
Core 20.5 15.5 11.5 20.5 15.0 11.4 20.8 15.0 11.3
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The summary of average changes in T2*, T1 and MTR values within both the ROIs and for both the loading steps are presented in Table 4. The statistical analyses revealed that the mean T2* values for the first step of loading, in the whole and the core sections, were decreased significantly by 13 ± 7% (P = 0.028) and 16 ± 5% (P = 0.017), respectively. For the second loading step, changes in T2* values were not statistically significant, but showed a decreasing trend (9 ± 2% (P = 0.059) and 7 ± 5% (P = 0.121) for whole and core sections, respectively).

Table 4.

The summary of changes (%) of T2*, T1 and MTR of all the scanned tendon samples at different loading steps.

T2* change1 T1 change MTR change
2KHZ 5KHZ 10 KHZ
1st load 2nd load 1st load 2nd load 1st load 2nd load 1st load 2nd load 1st load 2nd load
Whole section −13±7 (P=0.028)2 −9±2 (P=0.059) +10±11 (P=0.137) −10±14 (P=0.428) +3±4 (P=0.139) +2±6 (P=0.799 +2±5 (P=0.518) +3±3 (P=0.124) +10±18 (P=0.215 +2±5 (P=0.631)
Core section −16±5 (P=0.016) −7±5 (P=0.121) +8±12 (P=0.200) −8±15 (P=0.468) +4±4 (P=0.052) +2±4 (P=0.618) +1±5 (P=0.845) +5±5 (P=0.045) 2±6 (P=0.558) +3±4 (P=0.388)
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1

All values are presented as Mean±SD (%). Negative and positive values are representing decrease and increase, respectively.

2

P-value is calculated from two paired two tailed t student test

Mean T1 value variations were not statistically significant or consistent for the two loading steps, for either of the selected ROIs (Table 4). Although, the mean MTR values were increased slightly by loading within both selected ROIs, the differences were not statistically significant (Table 4). As summarized in Table 4, the T2* was the only studied biomarker that demonstrated significant changes by applying tensile loads consistently for all tendon samples (P < 0.05).

Figure 3 shows the T2* maps of three different tendon samples as examples (A-2, B-1 and B-2, respectively). The maps are presented at three different loading conditions; non-loaded, 15 and 30 N. The regions of higher T2* (indicated by green to red in the maps) decreased in size by static mechanical loading for all the samples. In other words, the blue regions of lower T2* values were expanding by the loading. Similar local variations in T2* values were observed on the maps of all examined tendon samples in this study.

Figure 3.

Figure 3

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The T2* map in three selected tendon samples, A-1, B-1 and B-2 at different loading conditions (non-loaded, 15 and 30 N). The regions of higher T2* values (indicated by green to red) of all samples decreased in size within the tendons by mechanical loading.

Figure 4 depicts the MTR maps of sample A-1, B-1 and B-2 (same samples in Fig.3) at different loading conditions. The loading resulted in an inhomogeneous distribution of MTR in sample A-1 with local maxima in the sites of lower T2* values (Fig.3). The MTR increase was obvious in few samples in this study such as sample B-1, however other example samples (such as sample A-1 and B-2) did not show clear MTR shifts or increase in the maps.

Figure 4.

Figure 4

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The magnetization transfer ratio (MTR) maps for different loading conditions (non-loaded, 15 and 30 N) at 2 kHz frequency offset and 1500° MT pulse preparation. The MTR map indicated higher values and non-homogenous distributions for loaded tendons particularly in sample B-1. Sample A-1 and B-2 showed less variations by loading.

4. Discussion

Human cadaveric peroneal tendons were studied using UTE-MRI under tensile static loading. This study follows prior works using UTE-MRI for assessing loaded tendons (Chang et al., 2015a, 2015c; Koff et al., 2014; Syha et al., 2014). Monitoring the UTE-MRI based biomarkers of tendons under tensile load application may provide an index for the tendon’s mechanical performance, which in turn can be used for early stage diagnosis of tendon diseases.

UTE-MTR and T2* were the two studied properties in the literature (Chang et al., 2015a, 2015c; Koff et al., 2014; Syha et al., 2014). UTE-MTR was found to be increasing in vivo after short-term exercises (Syha et al., 2014) and ex vivo under static tensile loads (5–10 N) (Chang et al., 2015a). Koff et al. observed T2* decreasing in tendons after cyclic loading (100 cycles of 45 N) (Koff et al., 2014). Nevertheless, Chang et al. (Chang et al., 2015c) did not observe any significant change in T2* of tendons under tensile static loads (10 N). This current study helps to realize the impacts of the load magnitude on the tendon MRI-based properties. It was illustrated that the level of changes in the UTE-MRI properties is a function of the applied load magnitude. Specifically, the level of changes in T2* was higher for 30 N load compared with 15 N load. It is likely that the 10 N load applied in a previous study (Chang et al., 2015c) was not enough to exert any changes in T2*.

The fabricated MRI-compatible loading device with plastic grippers provided a method to apply higher loads compared with our previous study (Chang et al., 2015c). Specifically, by eliminating the need for sutures, we were able to apply >10 N to the mounted tendons. To the authors’ knowledge, the applied loads on the peroneal tendons during daily activities have not been reported in the literature. However, for Achilles tendons, the estimated load peak, during recreational running, may reach up to several times of body weight (Andrew and Jonathan, 2014). This is much higher than applied loads in this study. The loading device employed plastic springs that ended in a very compact and sealed unit which was compatible with a wrist coil, without requiring strings and suspended loads (Chang et al., 2015c; Helmer et al., 2004; Wellen et al., 2005, 2004). However, the plastic spring may have a limited creep and stiffness reduction along the loading process in MRI. To reduce the stiffness changes along the loading experiment, springs underwent four loading sessions from 0 to maximum compression before scanning. Along the four loading sessions, the average calculated stiffness reduced up to 6% and converged to 1.98 N/mm.

Applying the first step static tensile load (15 N) resulted in significant reduction of mean T2* in all scanned tendon samples for both whole and core sections (13 ± 7% (P = 0.028) and 16 ± 5% (P = 0.017), respectively, Table 1 and 4). For the second loading step (30 N), the T2* values trended down, but this did not reach statistical significance (Table 1 and Table 4). One possible explanation is that the second loading step was performed after an hour of loading by 15 N. Such preconditioning may have prevented the tendons from demonstrating significant changes at the subsequent 30 N load application. T2* reductions were consistent with the observations from Koff et al. (Koff et al., 2014) in rabbit patellar tendon after cyclically loaded by 45 N. Therefore, both static and cyclic loadings resulted in similar changes in T2*. The applied load in our previous study (10 N) and the loading set up (suturing strips passing through the syringe and connecting to suspended weights), were likely limited in their ability to exert significant changes in tendons. Likewise, Helmer et al. (Helmer et al., 2004) did not observe changes in rabbit Achilles when 5 N tensile load was applied.

Of the biomarkers evaluated in our study, T2* was the only one that consistently demonstrated significant changes by applying mechanical loads for all tendon samples. Mean T1 value variations were not statistically significant for both of the loading steps and within both selected ROIs (Table 1 and Table 4). Mean MTR values were increased slightly by loading within both selected ROIs, however, the differences were not statistically significant (Table 3 and 4).

T2* maps demonstrated a shrinkage of high-T2* regions by the application of tensile load for all scanned tendons (Figure 3). High-T2* region shrinkage was observed in both core and rim sections of the tendon samples. However, more intense changes were found in the core sections (Figure 3). Such T2* reductions and shifts were expected due to, first, free water movement from denser regions of tendons towards less dense regions and possibly to both ends of tendon mounted in clamps, and second, reduction of crimp number as well as flattening of crimp angles. Specifically, such crimp reductions and flattening result in an angle decrease between collagen fiber axes and static magnetic field (B0), which in turn leads to lower T2* values. The lowest and highest values of T2* could be found at 0° and 55° angles with respect to B0. (Chang et al., 2015b; Mengiardi et al., 2006; Mountain et al., 2011). The T2* map shifts were not perfectly homogenous, probably due to uneven mechanical stress distribution in tendon cross sections (Slane and Thelen, 2014). Interestingly, the loading led to a slight increase of MTR in the sites of lower T2* values (Figures 3 and 4) which agrees with MTR increases, observed by Syha et al. (Syha et al., 2014) in vivo and by Chang et al. (Chang et al., 2015a) in vitro. Nevertheless, the MTR increase was obvious only in few samples of this study such as sample B-1 (Figure 4).

Despite the UTE-MRI techniques, the cilinical MRI has been used very often for loaded tendon studies. Helmer et al. (Helmer et al., 2004) investigated tendon proton density and to monitor intra-tendinous water movement after mechanical loading. Proton density demonstrated an increase in the rim sections of tendons while a decrease was seen in the core sections with 10 N loads, indicating outward water shifts (Helmer et al., 2004). Remarkably, for a lower tensile load (5 N), no change was observed in the core section. Wellen et al. (Wellen et al., 2005) studied T2 and T1 of tendons under loading. T2 increased in the rim but decreased in the core sections of tendon after loading. T1 decreased in both rim and core regions (Wellen et al., 2005). Repeated tensile loads after a period of unloading resulted in similar changes in T2 (Helmer et al., 2006). Mountain et al. (Mountain et al., 2011) also reported a significant T2 increase in tendons after static tensile loading, however the tendons were fixed in formalin after loading. They also examined the tendons in different orientation angles with respect to B0 direction. The T2 values were higher in loaded tendons for all orientation angles (Mountain et al., 2011). Apparent diffusion coefficient (ADC) was also invesitaed in other studies (Han et al., 2000; Wellen et al., 2004). A significant increase in ADC was observed and interpreted as a sign of water reduction in the tendons (Han et al., 2000), particularly in the tendon core (Wellen et al., 2004).

This study has several limitations. First, a small number of samples were included. As is the nature of pilot studies, our results require a validation with larger sampling sizes. Furthermore, our tendon specimens were obtained from a wide span in age of subjects and future studies should evaluate the possible effects of age. Second, the T2* calculations were performed using single-component exponential models. Future studies will include bi-component models to investigate whether the T2* reduction occurrs in the short or long T2* components of the tissue. It should be noted that the bi-component models require more data points with higher signal-to-noise ratio in order to ensure appropriate fitting. Third, the tendon samples were scanned only parallel to B0 direction. Although the T2* reductions can be explained by changes related to water movement, the reduction and flattening of crimps may also play a role. Future studies evaluating T2* may incorporate other tendon orientations, such as scanning at the magic angle (55°), for a comprehensive understanding of UTE-MRI changes of tendon under loading. Moreover, future studies may consider incorporating biomarkers that are known to be less sensitive to orientation angle, such as those generated from UTE-MT with two-pool modeling (Ma et al., 2017, 2016). Fourth, the employed plastic springs may demonstrate limited creep during the loading process and MRI scanning. Future studies should employ active MRI-compatible actuators (e.g., pneumatic actuators) to achieve higher accuracy and reliability. Finally, the degree of change in loaded healthy versus pathologic tendons would be of interest in future studies.

5. Conclusion

This study investigates the mechanical status of tendons through observations in UTE-MRI biomarkers. T2*, T1, and MTR values were measured for human peroneal tendons before and under static tensile mechanical loads. Among examined MRI properties, only T2* demonstrated significant changes under mechanical loading (over 12% reduction). T2* maps revealed a significant shrinkage of high-T2* regions in tendon cross sections which might be caused by, first, intra-tendinous shifts in water and second, alterations in tendon crimp. T2* monitoring over the course of loading can be suggested as a new technique for quantitative investigation of the tendons biomechanics.

Acknowledgments

The authors thank Niloofar Shojaeiadib for performing the statistical analyses. The authors acknowledge grant support from the VA (I01CX001388) and NIH (1R01 AR062581-01A1 and 1 R01 AR068987-01).

Footnotes

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Conflict of interest statement

The authors have no conflict of interests to disclose.

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