Mechanics of fresh, frozen-thawed and heated porcine liver tissue.

http://informahealthcare.com/hth ISSN: 0265-6736 (print), 1464-5157 (electronic) Int J Hyperthermia, 2014; 30(4): 271–283 ! 2014 Informa UK Ltd. DOI: 10.3109/02656736.2014.924161

ORIGINAL ARTICLE

Mechanics of fresh, frozen-thawed and heated porcine liver tissue Cora Wex, Anke Stoll, Marlen Fro¨hlich, Susann Arndt, & Hans Lippert

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Clinic for General, Visceral and Vascular Surgery, University Clinic Magdeburg, Germany

Abstract

Keywords

Purpose: For a better understanding of the effects of thermally altered soft tissue, the biothermomechanics of these tissues need to be studied. Without the knowledge of the underlying physical processes and the parameters that can be controlled clinically, thermal treatment of cancerous hepatic tissue or the preservation of liver grafts are based primarily on trial and error. Materials and methods: Thus, this study is concerned with the investigation of the influence of temperature on the rheological properties and the histological properties of porcine liver. Results: Heating previously cooled porcine liver tissue above 40 C leads to significant, irreversible stiffness changes observed in the amplitude sweep. The increase of the complex shear module of healthy porcine liver from room temperature to 70 C is approximately 9-fold. Comparing the temperatures 20 C and 20 C, no significant difference of the mechanical properties was observed. Furthermore, there is a strong relation between the mechanical and histological properties of the porcine liver. Temperatures above 40 C destroy the collagen matrix within the liver tissue. This results in the alteration of the biomechanical properties. The time-temperature superposition principle is applied to generate temperaturedependent shift factors that can be described by a two-part exponential function model with an inflection temperature of 45 C. Conclusions: Tumor ablation techniques such as heating or freezing have a significant influence on the histology of liver tissue. However, only for temperatures above body temperature an influence on the mechanical properties of hepatic tissues was noticeable. Freezing up to 20 C did not affect the liver mechanics.

Biothermomechanics, collagen, histology, shear rheology, soft tissue mechanics, tumor ablation

Introduction This work focuses on the changes in the mechanical properties of liver tissue as a function of temperature. One area of application still lacking basic research results is the thermal treatment of liver tumors. To date no generic rules for this kind of therapy exist, but the majority of patients with hepatocellular carcinoma are dependent on the thermal treatment of liver tumors as only a few patients are ideal candidates for surgical resection because of technical difficulties, age at the time of diagnosis and advanced cirrhosis [1]. The most common thermal technique is ablation, including both heating (radiofrequency, microwave, laser, and highintensity focused sonography) and freezing (cryoablation) [2]. Apart from high-intensity focused ultrasound, which can noninvasively treat tumors without injury to the surrounding tissue [3], ablation techniques have in common the minimally invasive or percutaneous insertion of a needle into the tumorous tissue. Ablation by heat then generates temperatures

Address for correspondence: A. Stoll, Clinic for General, Visceral and Vascular Surgery, University Clinic Magdeburg, Germany, Leipziger Straße 44, 39120 Magdeburg. Tel: 0049/391 6117-133. E-mail: [email protected]

History Received 1 August 2013 Revised 6 May 2014 Accepted 10 May 2014 Published online 19 June 2014

between 90 C and 110 C at the device tip which leads to temperatures between 60 C and 80 C in the surrounding tissue [1,2,4]. The necessary duration of the heat treatment depends on the ablation technique and the size of the tumor. The chance of an incomplete ablation increases with larger tumors [4]. The goal of thermal ablation using any of the existing systems is to kill the entire tumor and a circumferential cuff of normal liver surrounding the tumor to ensure an adequate tumor-free margin [4]. A recent study by Moffitt et al. [5] showed that the duration of a thermal treatment does not matter once a critical tissue temperature of 80 C was exceeded. Contrary to ablation by heat, cryoablation uses a freezing unit to generate temperatures of approximately 40 C with the help of liquid nitrogen. The freezing process is completed when the total tumor volume was encompassed by the ‘‘ice-ball’’ and when the thermocouples in the normal liver around the tumor revealed a temperature of 20 C or lower [6–8]. For ablation therapy, temperatures in the range of 60–90 C for heat ablation and temperatures equal to or below 20 C for cryoablation are applied to the liver tissue as these are the temperatures that ensure extensive damage of cancerous liver tissue. The temperatures in between are interesting for liver graft preservation as they are the temperatures needed to maintain cellular viability and organ function. Thus far, the major principle of organ preservation consists of reducing the


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metabolic activity by lowering the temperature. The metabolic rate shows a 12- to 13-fold decrease when the temperature is reduced from 37 C to 0 C [9]. Until now, liver grafts are flushed and cooled with specific preservation solutions at 4 C to reduce the metabolic activity of the biological tissue and to prevent cellular swelling [9–11]. However, recent studies show that the storage of liver with the warm perfusion technology at body temperature led to much better results [12]. However, it is well established that post mortem changes in the tissue due to biological degradation occur which inevitably also change the mechanics of liver tissue [13]. Stiffness is the most relevant mechanical parameter to describe changes in the mechanics of liver due to thermal treatment. Thermal lesions generated by heat treatment and tissue temperatures of 60–80 C have recently been measured to be 8–10 times stiffer than healthy liver [3,14,15], while several types of hepatic carcinomas have been measured to be 2–5 times stiffer than healthy liver under small strains [16–18]. After the heat ablation procedure, the size and position of the thermal lesion can be clearly identified as a stiffer region. Freeze-thaw cycles of liver tissue do not seem to have a significant influence on the stiffness of liver [19] but are clearly visible as a hypoechogenic area on the intraoperative ultrasonography. As the studies cited above, show that there are no generic results for the temperature dependence of the mechanical properties of liver tissue. Even though all studies agree on a general stiffness increase for heattreated liver, the results vary greatly and are hardly comparable. Mechanical changes of frozen and thawed liver are rarely reported on. Thus, the first objective of the present paper is to describe the mechanical behavior of thermally altered porcine liver. Tests were carried out using a parallel plate shear rheometer [13]. The presented dataset from 10 different temperatures and their influences on 8 healthy porcine livers was useful for the determination of mechanical changes of thermally altered liver tissue. In this paper we focused on the influence of the temperature. The effect of the duration of a thermal treatment will be investigated later on. Another objective is to use the rheological temperature dependent data in combination with the time-temperature superposition (TTS) principle. The proposed model describes the frequency and temperature dependent mechanical behavior of liver tissue in vitro. The TTS is often used to extend the mechanical characterization of polymer melts and polymer solutions over a limited frequency range. However, there have been few studies reporting the applicability of the TTS on soft tissues [20–22]. To our knowledge, the applicability of this method to liver tissue has not previously been shown. This paper indicates the suitability of the method. Thermally induced stiffness changes of the liver tissue must have their cause in the histological processes during heating/cooling. The major component of most soft tissues is protein. Thermal therapies lead to its heat-induced denaturation. Depending on the tissue type, hydration-level and tensile force on the tissue, critical temperatures for protein denaturation are found within a range of 40–50 C [3]. These temperatures will break the interactions in many proteins and denature them [23] and thus have an irreversible effect on the stiffness. Biologic cell death occurs due to the two processes

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called apoptosis and necrosis [14,24]. Thus, as another objective we want to determine the ability of mechanical testing to detect the pathological changes occurring in thermally altered liver tissue and to assess its accuracy in revealing the extent of histological tissue changes. Therefore, samples were histologically fixed and embedded in paraffin. Sections were Hematoxylin/Eosin stained and Gomori stained for light microscopy. The effect of the temperature on the histology of liver cells was investigated.

Methods Sample preparation Similar to previous investigations [13], we chose pigs from a heterogeneous pool of ecologically bred animals of various origins. Liver tissue was used in vitro and no IRB approval was required as the livers were obtained from a local butcher and qualified as food. The age of the pigs was approximately eight months and the weight was between 255 kg and 271 kg with the weight of the pig livers varying between 1.6 kg and 2.4 kg. Livers were stored in Krebs Ringer Hepes Buffer (glucose 10 g, L-Glutamine 100 mM, KCl 560 mM, Na2HPO4 2H2O 160 mM, MgCl2 6H2O 226 mM, CaCl2 2H2O 510 mM, NaCl 2.8 M, Hepes 100 mM) [25]. Samples of 35 mm in diameter and 3 mm in height were prepared. Further details on sample preparation are stated in a previous publication [13]. The samples were kept in Krebs Ringer Hepes Solution at 4 C until used. Only samples of isolated liver parenchyma excluding capsule were investigated. Samples did not contain any visible blood vessels. In order to investigate the influence of temperature on the mechanical properties of liver, we heated a beaker with Krebs Ringer Hepes Solution on a hot plate until the desired temperature was reached. Then the samples, placed on a flat sieve, were put in the heated solution. Direct contact of the samples with the bottom of the beaker and thus the hot plate was avoided. The internal sample temperature was observed by piercing an additional sample with a digital laboratory thermometer which was then discarded. Once the assigned temperature was held for 15 min, the samples were taken out of the heated solution and immediately put into Krebs Ringer Hepes Solution at room temperature for 5 min. A similar approach was followed in [26]. For the freeze-thaw process, the samples were gently padded dry, wrapped in cellophane and put in a freezer at 20 C. The samples were left there for at least 1 day and then taken out and thawed at room temperature. Testing protocol The rotary shear rheology experiments were carried out at room temperature (20 C) using a rheometer (Thermo Scientific HAAKE MARS) with a parallel-plate system, as previously described [13]. As a control, a 0.6 N normal force was applied to ensure contact between sample and upper plate. The testing protocol for each temperature included strain sweep, frequency sweep, relaxation test and controlled strain rate test. A new sample was taken out of the buffer solution for each test, gently padded dry and placed on the plates which have sandpaper (waterproof metal sandpaper,

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roughness 80) glued on to prevent the sample from slipping, as suggested by [27–29]. After each test, the sandpaper was replaced. In order to avoid large deviations due to the heterogeneity of the pigs, a full test matrix was performed on each liver. Each liver was investigated at all temperatures (20, 20, 30, 40, 45, 50, 55, 60, 70, 80 C) and all of the tests mentioned above were conducted at each temperature. This procedure was repeated for 8 different livers in order to generate 8 test results for each test and temperature.


Dynamic testing The parallel plate rotary system was used to determine the storage modulus G0 , the loss modulus G00 and the phase angle [13]. The complex shear modulus is equal to G ¼ G0 þ iG00 . It is considered difficult to obtain accurate soft tissue properties as to date there do not exist standardized measurement methods and biological soft-tissue is often characterized by large standard deviations for parameters such as elastic modulus, shear modulus or viscosity. One approach for the reduction of the large fluctuations is the introduction of preconditioning of soft tissue. However, similar to [30–32] the authors feel that preconditioning does not reflect the material’s actual properties in its natural state. As the authors wanted to change the physiological conditions of the liver tissue as little as possible when making their mechanical measurements, no preconditioning was applied, see also [13]. Nevertheless, the organ harvesting and sample preparations apply unnatural strains to the liver. Thus, the authors suggest the term minimal preconditioning instead of no preconditioning. Strain sweep Strain sweep oscillation tests were performed to determine the limit of the linear viscoelastic (LVE) region. A constant frequency of 1 Hz and varying strain amplitudes (0.0001–1) following a sinusoidal motion from low to high strain were applied [13]. Frequency sweep For a constant strain amplitude of 0.001, which for sample age 5 1 h is within the LVE region [13], the frequency was varied from 0.1–10 Hz with 6 measuring steps per decade and with 3 cycles for each frequency. Again, further details are stated in [13]. Relaxation test In order to observe the structural stability of the material with time, samples were held between the rheometer plates for 600 s with an applied stress amplitude of 1 Pa using rotary shear.

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distribution was carried out to obtain information on the flow behavior of the liver. Further details are stated in [13]. Histology For every temperature, histological samples were prepared from the parenchyma before and after mechanical testing (after the rotational test). Samples were histologically fixed in formalin and embedded in paraffin. Sections were Hematoxylin/Eosin stained and Gomori stained for light microscopy. An Olympus BH-2 microscope (Olympus Deutschland GmbH, Hamburg, Germany) was used.

Modeling The time-temperature superposition principle is applicable when data can be shifted to and from a reference temperature T0 to form a master curve [20,21,33]. When no smooth master curve can be obtained, the TTS principle is not applicable. The master curve at the reference temperature T0 is equal to: G ð!, TÞ ¼ G ðaT !, T0 Þ,

ð1Þ

with the horizontal shift factor aT .

Results LVE limit An earlier investigation of the influence of the preservation time on mechanical parameters of the liver at room temperature showed an LVE limit of 0.8% for preservation times smaller than 1 h. For preservation times higher than 1 h, no clear LVE behavior was observed [13]. The absence of a clear strain independent region was confirmed in this study for all investigated temperatures as for logistical reasons post mortem times were slightly above 1 h. However, as the slopes of G0 and G00 in Figure 1(a and b) are very small or even close to zero for temperatures larger than 55 C, a quasi shear strain independent region for shear strains up to 0.5% is assumed which shows a deviation of less than 5% of the mean value within the quasi LVE region. Mean G0 and G00 are the smallest for the frozen-thawed liver samples. Comparing room temperature and 45 C for a deformation of 0.1% shows an increase of 65% and 52% for mean G0 and mean G00 , respectively. The increase is slightly less pronounced at 1% deformation with 55% and 44% for mean G0 and mean G00 , respectively. The temperatures 70 C and 80 C lead to near constant mean G0 and mean G00 , Figure 1(a, b and d). Comparing room temperature and 70 C for a deformation of 0.1% shows an increase of mean G0 and G00 by a factor of 9.3 and 7.8, respectively. For a deformation of 1%, G0 and G00 increase by a factor of 9.1 and 7.6, respectively. The parameter tanðÞ decreases with increasing temperature for shear deformations between 0.04% and 5%, Figure 1(c).

Rotational tests Rotational rheometry was conducted to find structural and compositional changes of the investigated material and to measure shear viscosity. A controlled strain rate test with _ varying from 0.0001 to 0.01 s1 and a logarithmic step size

Frequency dependence The influence of the preparation temperature on the frequency dependence of mean G0 and G00 is shown in Figure 2(a and b), while Figure 2(c) presents the temperature and frequency


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Figure 1. Storage modulus (a), loss modulus (b), tanðÞ (c) dependent on shear strain, storage and loss modulus and tanðÞ dependent on temperature for ¼ 0.001 and ¼ 0.01.

dependence of the dynamic viscosity . Mean G0 , G00 and for room temperature are equivalent to storage and loss modulus and viscosity found for 20 C for all investigated frequencies. Mean G0 and G00 increase with increasing frequencies. Comparing for example f ¼ 0.1 Hz and f ¼ 10 Hz at 20 C shows an increase of mean G0 by a factor of 2.1. For 80 C, this increase is less pronounced with a factor of 1.7. Mean G00 increases by a factor of 2.9 at 20 C and by 1.8 at 80 C comparing the frequencies f ¼ 0.1 Hz and f ¼ 10 Hz. In contrast, the mean dynamic viscosity decreases significantly with increasing frequency. Comparing f ¼ 0.1 Hz and f ¼ 10 Hz at 20 C shows a decrease of mean by a factor of 47.3. At preparation temperature 80 C, the mean dynamic viscosity decreases by a factor of 58.2 comparing f ¼ 0.1 Hz and f ¼ 10 Hz. Increasing the preparation temperature from 20 C to 70 C increases mean G0 by a factor of 10.4 for 0.1 Hz, of 9.4 for 1 Hz and of 9.1 for 10 Hz. Mean G00 increases by a factor of 9.9 for 0.1 Hz, of 8.4 for 1 Hz and of 6.4 for 10 Hz. The mean dynamic viscosity increases by a factor of 10.3 for 0.1 Hz, of 9.3 for 1 Hz and of 9.1 for 10 Hz. Preparation temperatures higher than 70 C do not change mean G0 , G00 and any further. Mean G0 is larger than G00 for all temperatures.

The magnitude of the complex modulus jG*j is shown in Figure 3. The quotient of mean G0 and mean G00 which is equal to mean tanðÞ was investigated to reveal the ratio of the viscous and the elastic portion of the deformation behavior [33]. Figure 4 shows that mean tanðÞ fluctuated between 0.2 and 0.35. The mean value of tanðÞ gives similar results for f ¼ 0.1 Hz and f ¼ 1 Hz, while f ¼ 10 Hz leads to higher mean tanðÞ values. For all investigated frequencies mean tanðÞ at 20 C is smaller than at room temperature. For increasing temperatures equal or larger than 20 C mean tanðÞ also decreases. This effect is more pronounced for f ¼ 10 Hz. As Figure 5 shows, the complex shear modulus jG*j can be shifted along the horizontal frequency axis to obtain a smooth master curve at a reference temperature of 20 C. For better readability of Figure 5, the diagram is limited to the master curves of the first 5 specimen (instead of all 8). The curves of jG*j show well overlapping areas and the frequency domain could be extended to a significantly larger number of decades. The mean horizontal shift factors aT ðTÞ for each temperature and all 8 specimen were calculated and are shown in Figure 6. Clearly, the shift factor at the reference temperature is 1. The points below 45 C lie on a straight line. The points above 45 C lie on a steeper straight line. The application of a linear



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Figure 2. Storage modulus (a), loss modulus (b), and viscosity (c) dependent on frequency.

Figure 3. Mean complex modulus jG*j compared to results by the references [14] and [26].

regression approach in Matlab gives the following fits of exponential functions: Below 45 C: aT ¼ 5:0 102 e0:15T :

ð2Þ

aT ¼ 2:7 107 e0:42T :

ð3Þ

Above 45 C:

Figure 4. Mean tanðÞ compared with reference [26].

Figure 6 shows the fit of the two exponential functions. The intersection of the two functions at T ¼ 44.9 C can be called the inflection point and generates a biphasic plot. Flow behavior Within the strain rate interval of 0.0001 – 0.01 s1 the viscosity is nearly constant and the shear stress has a linear appearance in the loglog plot in Figure 7. The standard deviation for and are only plotted exemplarily for 20 C


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Figure 5. Master curves of jG*j shown for 5 specimen.


Figure 7. Flow curves temperatures.

and

viscosity

for

varying

preparation

Figure 6. Mean of experimentally obtained shift factor and curve fit with two exponential equations.

and 80 C for reasons of clarity and readability of the diagram. At the maximum strain rate of 0.01 s1 a peak shear level of 16% was reached for 20 C. For increasing temperatures the peak shear level decreases down to 10% for 80 C. The temperature dependence of the shear stress and the viscosity is shown in Figure 8. For low preparation temperatures, shear stress and viscosity are almost independent of the temperature. Especially for low strain rates, the shear stress is independent from preparation temperatures up to 40 C. For higher strain rates, the shear stress independence ends at 20 C. The temperature independent region is then followed by an increase of and up to 70 C, which gives the maximum values for and . Comparing room temperature and the preparation temperature 70 C, the parameter increases by 9.7 for _ ¼ 0.0001s1, by 11.9 for _ ¼ 0.001 s1 and by 9.6 for _ ¼ 0.01 s1. Similarly, the parameter increases by a factor of 10.0 for _ ¼ 0.0001 s1, by 12.9 for _ ¼ 0.001 s1 and by a factor of 9.3 for _ ¼ 0.01 s1. Histology At 20 C preparation temperature, normal liver parenchyma that is arranged in lobules can be observed. Within the

Figure 8. Shear stress and viscosity with increasing temperature and _ ¼ 0.0001, 0.001 and 0.01 s1.

lobules the hepatocytes are arranged in cords separated by sinusoids. Sinusoids are vascular channels lined by the endothelium [34]. Bile canaliculi as a network of tiny passages are contained within each cord. For samples prepared at 20 C, the HE stain shows stable connections between the hepatocytes, which are unaffected by mechanical testing, see Figure 9(a and b). The Gomori stain shows intact reticular fibers located in the space of Disse before and after mechanical testing (Figure 9c and d). For preparation temperature 80 C, the arrangement of cells and cords is destroyed. Sinusoidal dilatation and bile canaliculi dilatation become visible (Figure 10a and b). The destruction is largest around the central vein of each lobule. However, the structure of the lobules with their surrounding connective tissue is still intact. In Figure 10(c and d), the Gomori stain shows a destruction of the reticular fibers. During the denaturation process cells within the fiber lattice break down. With the supporting structure of the reticular fibers missing, the

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Figure 9. Liver parenchyma at preparation temperature 20 C, (a) HE before mechanical testing, (b) HE after mechanical testing, (c) Gomori before mechanical testing, (d) Gomori after mechanical testing.


hepatocytes separate. The disintegration of hepatocytes is the most dominant in the center of each lobule where the central vein is located. In this area the sinusoids are larger. With mechanical testing, sinusoidal dilatation and bile canaliculi dilatation increase. The cell structure of the samples that were frozen at 20 C (Figure 11) and thawed is damaged due to the formation of ice crystals during the freezing process. Sinusoidal and bile canaliculi dilatation can be observed. However, the cord structure of the hepatocytes is still intact. The Gomori stain shows that the reticular fibers are also still intact. The liver histology at 20 C is largely unaffected by mechanical testing.

Discussion Experimental procedure One purpose of this paper was to collect data from rheological shear tests to determine the effects of the preparation temperature on the mechanical properties of porcine liver parenchyma. For a better understanding of the processes occurring when thermally altering liver tissue, the biothermomechanics of porcine liver tissue were studied. However, several issues need to be addressed before results can be interpreted. In the following, the authors point out the simplifications that were necessary to conduct the rheological experiments:

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(i) Experiments were conducted on porcine tissue. Even though there are apparent differences between human and porcine tissue concerning the capsule and capsule failure characteristics [35,36], the overall tissue and cellular distribution patterns of porcine liver tissue are similar or identical with those in humans [37,38]. Many similarities were found to be supporting the suitability of the pig as a promising substitute to study the liver pathology [37]. Porcine livers are even used for bridging which means the use of porcine livers for temporary transplantation in humans [38,39]. (ii) The mechanical properties of liver were measured in vitro. Ex vivo and especially in vitro testing conditions can be quite different from the in vivo situation (perfusion, body temperature, physiology). Mazza et al. [17] found the ex vivo human liver mean stiffness 17% higher than the in vivo liver. The stiffness of ex vivo unperfused lobes was found to be 50% higher than for the in vivo perfused liver, but a stiffness increase of only 17% was observed between in vivo and ex vivo perfused tests [40]. Relaxation tests showed that the amount of relaxation in the tissues appeared to be greater post mortem than in vivo [41]. Hines-Peralta et al. [42] found that larger zones of coagulation were achieved for in vivo liver than for ex vivo liver with short energy applications. This finding is in contradiction to virtually all prior reports concerning RF and microwave ablation, where in vivo results have been substantially smaller than those achieved ex vivo, a fact that is largely attributed to the negative effects of blood perfusion [42]. Because of the apparent differences between ex vivo and in vivo material properties of porcine liver and discrepancies in the literature, the prediction of in vivo heatinduced changes remains problematic. (iii) The thermal treatment was applied to healthy liver tissue instead of tumorous tissue. This should be kept in mind when comparing these results with studies on ablation of tumorous tissue. For example, [17] state that the stiffness of tumorous liver tissue is approximately 3 times higher

than that of healthy tissue. Thus, the initial stiffness of tumorous tissue is already elevated. (iv) Tests were conducted on tissue samples rather than on a local area of an entire organ, and the thermal treatment was applied to change the mechanical properties of the entire sample and not just a local area which is typical for the use of a thermal probe in the ablation process. (v) Livers were stored in Krebs Ringer Hepes Buffer which is suitable for preservation of soft tissue. However, currently the use of UW solution [11] or Custodiol [43] is more common in graft preservation. Furthermore, there are some experimental issues that have to be addressed here: (i) Similar to a wide range of soft tissue experiments [13,28,31,44], our tests were conducted using a parallel plate rheometer. This geometry allows to measure threedimensional structures, such as soft solids [33] even though there is no constant shear gradient in the gap between the plates. For the measurement of linear viscoelasticity a uniform shear rate is not required. (ii) Dehydration was found to play a minor role in our testing conditions. In preliminary investigations, stress relaxation tests were conducted to investigate the dehydration behavior of the liver tissue. Dehydration was negligible within test duration [13]. (iii) Slippage of the samples was avoided by the use of sandpaper on both plates [13]. (iv) The heating of samples was performed in a water bath to generate consistent temperatures within the entire sample. This method is intended as an approximation of thermal ablation. (v) There is a difference between the stiffness of liver parenchyma with and without capsule [30,44–47]. Seki and Iwamoto [45] conducted compression tests to determine the breaking stress of ex vivo liver which was 22% lower with removed capsule than with intact one. A similar result was obtained by Brown [30]. An even



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more severe influence of the capsule on the elasticity of the liver parenchyma was observed by Carter et al. [47] who conducted indentation tests and showed that the elasticity with capsule is 2 times larger than without. Another approach used for testing adult porcine and bovine liver samples with a torsional resonator device was followed [44,46]. For this measuring principle it was found that the presence of the organ capsule leads to a considerable increase of the shear stiffness of the liver tissue. For frequencies from 1.3 to 12 kHz the dynamic modulus jG*j of capsule was estimated almost one order of magnitude larger than jG*j of the internal parenchyma. Thus, for stresses applied normal to the liver surface, the presence of the capsule has a significant effect on the overall organ stiffness as its stiffness dominates the response compared with the softer parenchyma. But this finding seems to depend entirely on the measurement technique as we did preliminary shear rheology tests on liver samples with and without capsule and no significant differences were observed. The capsule has a negligible thickness and a significantly higher stiffness than the parenchyma. Therefore it would contribute little to the observed shear deformation. This leads to the conclusion that our shear rheology experiments are not useful for detecting differences of liver parenchyma and capsule. Nevertheless, it is a suitable method to obtain information about the mechanics of liver parenchyma and its changes with varying temperatures.

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equal to 45 C [3]. Contrary to our findings, [52] states that the post mortem mechanical properties of bovine liver do not change for experiments conducted at room temperature or body temperature. No statistical difference in the failure stress and strain between the two temperatures was found. The decreasing tanðÞ for increasing preparation temperature in Figure 1(c) shows that the heated samples become slightly less viscous than samples at room temperature. Figure 1(d) shows that there are two plateaus for the temperature dependence of viscosity (20–30 C and 70–80 C) that cannot be described by the temperaturedependent continuous change of the viscosity that is typical for a single collagen fiber as introduced by Arrhenius [53]. However, the course of the viscosity curve with its two plateaus agreed well with the general course of the viscosity curve obtained by Zoldak et al. [54]. They investigated the protein glucose oxidase from Aspergullus niger. It shows a similar viscosity course for temperatures from 20–80 C with plateaus at 20–40 C and 60–80 C. The plateau observed for temperatures 60–80 C hints towards the fact that a further increase of the preparation temperature does not change the mechanical properties of the liver any further. Again, a oneway ANOVA with a significance level of 5% showed that there are no significant differences in mean G0 and G00 for 60 C, 70 C and 80 C. All changes in the mechanical parameters observed in this study are irreversible as they are obtained after the heat treatment when the tissue temperature is brought back to 20 C and thus are considered permanent. Temperature and frequency

LVE behavior 0

00

There is a small downward slope in mean G and G for the temperatures 20, 20, 30, 40, 45, 50, 55 C. For 60, 70 and 80 C the mean G0 and G00 follow a slight arch. This shows that there is no clear strain independent region. However, since the change in mean G0 and G00 is so small a quasi strain independent region for shear strains of 0.1–1% can be assumed with a standard deviation of less than 10% for all temperatures. Similar to [19,48–50], we found that the freezing-thawing process does only minimally influence the mechanical parameters even though the formation of ice crystals leads to irreversible damage of the tissue [51]. For increasing preparation temperatures, the mean G0 and G00 increase. Figure 1(a and b) indicate that for an increase from room temperature to the porcine body temperature of approximately 40 C mean jG*j increased by 164% for a shear strain of 0.1% and by 152% for a shear strain of 1%. Mean tanðÞ decreased by 8%. Thus, a post mortem temperature increases the liver up to body temperature after extensive cooling at 4 C and changes its mechanical properties slightly. Conducting a one-way ANOVA with a significance level of 5% shows that there is no significant difference between the mean G0 and G00 for 20, 20 C and 30 C. However, the increase in mean G0 and G00 for 40 C and higher is considered significant, assuming a significance level of 5%. This confirms ultrasound measurements of the stiffness of bovine liver suggesting irreversible tissue changes for temperatures of greater than or

The mean G0 and G00 are nearly identical for samples kept at room temperature and for samples frozen and thawed – as is the viscosity. For temperatures higher than 20 C, the shear moduli and the viscosity increase up until the preparation temperature reaches 70 C. For temperatures of 70 C and higher, the shear moduli and viscosity are independent from the preparation temperature. This indicates, that the ablation process could be conducted with tissue temperatures as low as 70 C. Mean G0 is larger than G00 for all temperatures. This indicates a stable structure and a dominating elastic behavior for all temperatures. The results obtained from the frequency sweep are generally in agreement with results stated by [14] and [26] where dynamic mechanical analysis is applied to study canine and porcine liver mechanics, respectively. Figure 3 shows that mean jG*j is of the same order of magnitude throughout, but the steep increase of jG*j was observed for slightly higher temperatures by [14] and [26], namely between 60 C and 70 C, while the current results showed a more gradual increase between 40 C and 70 C. This difference might be based on the different sample sizes and on the fact that Kiss et al. [26] and Bharat et al. [14] conducted compression tests normal to the sample surface while ours were shear experiments, and liver is known to be anisotropic (not equal in all directions) [55]. Additionally, samples were stored in distilled water in [14] and in isotonic saline in [26] instead of in a solution which ensures good preservation of the tissue. Furthermore, no information about the preservation times of


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their samples was given. Another aspect that might explain the observed differences of the frequency and temperature dependence of the liver is that researchers in [26] froze their samples after the heat treatment and thawed them for the dynamic mechanical measurements. The overall increase of mean G0 and mean G00 from 20 to 70 C is approximately 10-fold and 7-fold, respectively. This agrees well with Sapin-de Brosses et al. [3] who conducted ultrasound experiments on bovine liver and observed a shear modulus increase of 10 times for a temperature increase from 45 to 67 C. Bharat et al. [14] observed an increase of the Young’s modulus from temperature 60 C to 90 C which was 5-fold while Kiss et al. [26] observed a 7-fold increase of the complex modulus for a temperature increase from 60 to 90 C. Kiss et al. [26] and Bharat et al. [14] also detected a local minimum of jG*j at 85 C and 70 C, respectively, which was not observed in our experiments. Since Kiss and Bharat do not state the number of experiments conducted for each preparation temperature and only give incomplete information about the statistical basic data, no statement about the significance of their results can be made. It is also important to keep in mind that we actually measured mean jG*j from testing in shear while Kiss and Bharat obtained a mean jE*j from dynamic mechanical testing in compression. A Poisson’s ratio typical for soft tissue of 0.5 was assumed to convert jE*j into jG*j; see for example [56]. The considerable increase of mean jG*j with increasing frequency is consistent with [26]. The parameter tanðÞ was smaller in the current experiments and varied between 0.2 and 0.3 while in [26] values between 0.3 and 0.4 are stated. Thus, these results indicate the liver to be slightly more elastic and less viscous. The same reasons mentioned above apply here to justify the difference in the results.

Dewey et al. [58] who established a ‘‘break’’ temperature of approximately 43 C many years ago. The approximate protein content of raw porcine liver is 21% (71% water, 4% lipids, 2% carbohydrates, minerals, vitamins) [59]. When liver cells denature, bonds break and protein molecules unwind. Then the connective tissue shrinks as water is squeezed out and the protein molecules coagulate. The hyperthermic unfolding and subsequent aggregation of proteins leads to cell death. Among few others, Lepock et al. [60] investigated protein denaturation of hepatocytes. Different transition temperatures were noticed applying differential scanning calorimetry. No cell killing was observed for temperatures up to 43 C and 50% killing for 46 C [60]. Thus, temperatures between 43 and 46 C mark a transition from one temperature process into another, which coincides with our inflection temperature of approximately 45 C. However, the significant increase in cell death affecting the changes in the liver mechanics with increasing temperature remains unclear. The effect of protein denaturation on both cells and cell structures such as tissues will have to be investigated further in the future. A model, completely determined by thermodynamics and the biological response to heat induced changes in protein-folding has not been derived yet but would be an important contribution for an improved definition and quantification of thermal dose [61]. However, a first step towards a comprehensive theoretical description of heat induced protein denaturation was done with the introduction of the TTS principle. The resulting exponential equations allow the calculation of the viscoelastic properties of liver tissue at various temperatures over an experimentally convenient time or frequency range. The application of the Arrhenius or WLF equation was not sufficient to describe the experimental results.

The TTS principle

Temperature dependence of the flow behavior

Similar to references [20] and [21], the authors applied the TTS principle to temperature and frequency dependent soft tissue data and proved its applicability to liver tissue. Mechanical properties measured at temperatures ranging from 20 C to 80 C can be shifted to a reference temperature of 20 C by applying the TTS. The experimental shift factors cannot be fit by one straight line, which means that the underlying thermal mechanism is very complex. It appeared that two separate exponential functions were most suited to fit the horizontal shift factors for 20–40 C and 45– 70 C obtained from the rheological experiments. The temperatures 20 C and 80 C were not included in the model as on both ends of the investigated temperature range the mechanical parameters investigated had become temperature independent. The two model functions result in an inflection temperature of 45 C which most likely represents the transition from one degrading temperature process into another. For temperatures lower than 45 C the mechanical properties of porcine liver changed slowly due to heating while for temperatures higher than 45 C the changes in the mechanical parameters occur faster with an increase of the model exponent from 0.15 to 0.42. This critical temperature obtained for porcine liver tissue confirms recent studies by Sapin-de Brosses et al. [3], Sapareto et al. [57] and

Relaxation tests conducted by [28] indicated that liver is a fluid-like or very soft viscoelastic material. This was confirmed by our controlled strain rate test results. The nearly constant viscosity indicates Newtonian-like properties for strain rates between 0.0001 s1 and 0.01 s1. This means, that the flow behavior of liver is dominated by its viscous properties. As the applied strain rates are quite small, we might only measure the linear region of the stress-strain curve as mentioned in [41] or [62]. The shear stress – strain rate relation for a Newtonian fluid can be described by ¼ _ with a constant value of for each investigated temperature. However, it is a fact that the liver is a viscoelastic material even though conducting rotational rheometry with a controlled strain rate test showed Newtonian behavior. Apparently, the Cox-Merz rule, which has been used for polymer melts, is not applicable to cross-linked systems such as soft tissues [63]. Thus, we cannot assume the steady shear viscosity to be equal to the dynamic viscosity. In order to define the relation between those two parameters and to learn more about the change from viscous to viscoelastic behavior, further investigations will be conducted in the future. For strain rates higher than _ ¼ 0.05 s1, and thus shear levels between ¼ 40% and ¼ 80%, the samples started to

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buckle and the measurements became unreliable. This happened because the sample was firmly connected to the upper and lower rheometer plates and the sample material could not accommodate any more shear and started to buckle. Thus our observations are limited to this small strain rate interval. The viscosity is almost constant for the investigated strain rate interval but dependent on the preparation temperature. This flow behavior also occurs for polymer melts. The highest viscosity of approximately 4 104 Pas was obtained for samples prepared at 70 C which is very high and comparable with natural rubber or polymer melts [33]. An increased viscosity and an increased shear stress for high preparation temperatures is equivalent to an increased stiffness. Histology Sinusoidal dilatation and bile canaliculi dilatation appear to be the severe histological changes in the liver cell structure for preparation temperatures 20 and 70–80 C. Additionally, only at temperatures above 40 C a fracture in the reticular fibers was observed. The denaturation of collagen leads to the breaking of hydrogen bonds and an irreversible transformation of the cross-linked triple helical structure into a more random, coiled structure [64,65]. This led to the assumption that not the sinusoidal and bile canaliculi dilatation is responsible for the mechanical changes observed in this study but the state of the reticular fibers. Reticular fibers are composed of collagen. The fact that collagen submitted to cryopreservation does not suffer any histological modifications was shown in a study[66]. Thus, it can be assumed that the collagen fibers of the liver tissue do not undergo histological changes in the freezing process as they seem intact in the histological images of the frozen-thawed liver samples in Figure 10(c and d). We assume this to be the main reason for not finding significant mechanical differences between samples prepared at 20 C and frozen-thawed samples. Investigations on heated collagenous connective tissue in raw beef muscle samples showed that a medium to large proportion of granular connective tissue was noted compared to mostly wavy or straight fibers in unheated tissue [67]. Thus, the denaturation of the connective tissue is accompanied by the granulation and disintegration of collagen which leads to significant mechanical changes of the supporting tissue as described in this manuscript. The dilatation of the sinusoids and the bile canaliculi, which was observed at 20 C and 80 C, does not seem to have a strong influence on the mechanics of the liver samples.

Conclusions Our investigations lead to the following conclusions: (i) Heating previously cooled porcine liver tissue above 40 C leads to significant, irreversible stiffness changes observed in the amplitude sweep. (ii) All conducted tests showed no significant difference of the mechanical properties for the temperatures 20 C and 20 C. Thus, freezing and thawing only minimally affects the mechanical properties investigated in this study.


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(iii) For temperatures higher than 40 C denaturation of the tissue is assumed to take place. Temperatures between 70 C and 80 C did not seem to influence the mechanical properties of the liver any further as constant viscoelastic parameters were observed. Thus, tissue temperatures as low as 70 C are considered safe in ablation procedures. (iv) For the temperature range examined, the TTS principle is applicable to liver tissue. The two-part exponential function model with an inflection temperature of 45 C gave a good fit. (v) Two near constant plateaus for the investigated parameters were observed in every test which do not follow the Arrhenius equation typical for one collagen fiber [53]. (vi) With the shear modulus being a measure of the stiffness of the liver, a mean increase in G* of a factor of 9 was observed in the amplitude and frequency test confirming earlier studies by [3,14,15]. The initially lower stiffness of the investigated healthy tissue compared to cancerous tissue with an already elevated stiffness most likely did not affect the general increase of the stiffness with increasing temperature. (vii) There is a strong relation between the mechanical and histological properties of the porcine liver. The mechanical properties are strongly correlated with the state of the reticular fibers in the tissue. Temperatures above the body temperature of pigs change the histology of liver tissue significantly, leading to the destruction of the collagen matrix for temperatures of approximately 45 C and higher, which results in the alteration of the biomechanical properties of liver tissue. A correlation between the dilatation of the sinusoids and the bile canaliculi and the changes in the liver mechanics was not found.

Acknowledgements The authors wish to thank Eberhard Pietsch for helpful discussions and suggestions and Laura Bisberg for proofreading the manuscript.

Declaration of interest The authors report no conflicts of interest. The authors alone are responsible for the content and writing of the paper. This research is supported financially by Saxony Anhalt, Germany.

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