Name: a coupled experiment-finite element modeling methodology for assessing high strain rate mechanical response of soft biomaterials
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The authors would like to recognize the Center for Advanced Vehicular Systems (CAVS) and the Agricultural and Biological Engineering Department at Mississippi State University for supporting this work. This material is based upon work supported by the U.S. Army TACOM Life Cycle Command under Contract No. W56HZV-08-C-0236, through a subcontract with Mississippi State University, and was performed for the Simulation Based Reliability and Safety (SimBRS) research program. Also, this material is based upon work supported by the National Nuclear Security Administration, (Department of Energy) under award number [DE-FC26-06NT42755]. Finally, the authors would like to thank Mr. David Adams, Mr. Michael McCollum and Ms. Erin Colebeck for their effort in this research.

NOTE: Ethics Statement: The current work is unique to the institution&#8217;s research policy, and strictly follows the appropriate bio-safety and Office of Regulatory Compliance (ORC) guidelines.
1. Biomaterial Specimen Procurement
<ol>
	<li>Wear personal protective equipment in accordance with standard biosafety protocols of the laboratory and/or institution. Wear closed-toed shoes, long pants, a lab coat, surgical gloves, a protective mask, and safety goggles while handling porcine tissue...

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The reported methodology that couples the SHPB experiment and FE modeling of the SHPB offers a novel and unique technique to assess the uniaxial true stress-strain response of a biomaterial at high strain rates. In order to procure mechanical properties intrinsic to the native tissue, care must be taken to keep the biomaterial specimen between 5.56&#8211;7.22 &#176;C before SHPB testing. If the specimen is cooled below 5.56 &#176;C, water present in the tissue starts to crystalize into ice and subsequently changes the ti...

Motivation
The cardinal goal of the coupled Split &#8211; Hopkinson Pressure Bar (SHPB) experiment/finite element modeling of soft biomaterials (such as brain, liver, tendon, fat, etc.) was to extract their uniaxial mechanical behaviors for further implementation in human body FE simulations under injurious mechanical loads. The human body Finite Element (FE) model consists of a detailed human body mesh and a history dependent multiscale viscoelastic-viscoplastic Internal State Variable (ISV) material model for various human organs. This human body model can be used for a framework to build better standards for injury protection, to design innovative protective gear, and to enable occupant centric vehicular design.
Two modes of high rate injury have been widely observed in human trauma: explosive blast and blunt impact. Blast damage from explosive weaponry is the primary source of traumatic injury (TI) and the leading cause of death on the battlefield1. When detonated, these explosives form an outward propagating shock wave that produces large and abrupt accelerations and deformations. The resulting loads pose serious threats to those exposed. Although any part of the anatomy can be injured by shock waves, the prime areas of concern are (1) the lower extremity due to its close proximity to the ground, and (2) the head since injuries can inhibit normal brain function and survival2,3. These injuries can be categorized as primary, secondary, or tertiary injuries depending on the type of injury sustained. Because the strength of an explosive is characterized by its weight or size, standoff distance, positive pulse duration, and medium through which it travels, it can be difficult to adequately categorize these injuries3-6. Congressional reports indicate that military personnel have suffered nearly 179,000 traumatic injuries due to explosive weaponry and vehicle crashes in Iraq and Afghanistan from 2000 through March 20102. Due to the nature and locations of modern combat, head injuries are a leading concern for both the military and civilians3.
Aside from combat scenarios, TI has a variety of causes including automotive trauma; rodeo, motorcycle and domestic accidents; and sports injuries. For instance, despite improvements to safety equipment and protocols, mechanically induced traumatic brain injury (TBI) continues to be a leading source of mortality and lifelong morbidity in the U.S. The Center for Disease Control and Prevention (CDC) reports approximately 1.4 million TBI events each year, of which nearly 50,000 are fatal. American football alone accounts for more than 300,000 TBIs each year7. Survivors of such injuries are at risk for long-term neurological complications related to sensation, cognition, and communication. At this time there are approximately 5.3 million Americans living with these chronic disadvantages and disabilities. Direct and indirect U.S. medical costs from 2000 to 2010 totaled $60 billion8. However, these numbers do not account for non-medical costs and losses, or those incurred by the families and friends supporting TBI patients. Beyond purely economic analysis, TBI-induced disability creates a significant reduction in quality of life that can manifest as a significant burden on families and society.
The need for further understanding of the formation, characterization, and prevention of TI is clear. Biomechanical studies of the underlying mechanisms that cause TI provide insight and opportunity to reduce exposure or improve safety features for those at potential risk for TI. Furthermore, more advancement of the general understanding of TI formation may improve diagnostic methods and criteria, providing medical professionals who treat TI with better means of improving outcomes and saving lives.
A better knowledge of injury mechanisms and a better understanding of the biomechanics of injury development are needed to develop effective protective measures for the human body. Historically, simulations aimed at predicting injuries have been hampered by computational restrictions as well as the fidelity of the anatomical and material models employed. Full body simulations have focused on the overall loads on each body part, but the local stress, strain, and damage in each organ, muscle, bone, etc. has not been observed. For example, shoulder moment models use the dimensions of the arm, the load, and the applied angle to search for tabular values that specify whether or not a particular scenario is hazardous. A calculation of that nature is helpful for quick estimates but cannot capture what is happening locally from the hand all the way to the shoulder, especially when damage and injury are intrinsically local. Secondly, FE simulations have been used to capture the local response. The limitation in these efforts has not been FEA itself, but the material models that define each body part&#8217;s behavior under blast injury loads. Previously employed material models are adapted from simpler materials and have not endeavored to capture the myriad of complex mechanical behaviors exhibited by biological tissues. Therefore, high-fidelity computational models with ISV material models for organs in the human body represent the most realistic way to investigate the physics and biomechanics of TIs, to design innovative protective gear, and to establish better standards for injury metrics.
Background on Split-Hopkinson Pressure Bar (SHPB) and Internal State Variable (ISV) Material Model
Due to ethical issues involved with in&#160;vivo testing of human organs and the logistical issues associated with broad-scale human cadaveric testing, the current research effort involves mechanical experiments in&#160;vitro using specimens prepared from organs extracted from animal surrogates (e.g., pig as a most frequently used surrogate). Polymeric SHPB has been the preferred method for in-vitro testing soft biomaterials at high strain rates. The relevant deformational behaviors from SHPB testing and corresponding tissue damage-related information from the microstructural features of the tissue are incorporated into our ISV material models for organ mechanical descriptions9-10. These material models are then implemented into our virtual human body model to conduct FEA of various injuries. This process enables us to move towards the goal of accurately predicting the physics and nature of an injury for a given organ under diverse mechanical loading conditions (e.g. blast-induced, car crash and blunt impact) without the need for further physical experimentation. In order to accurately describe the phenomenological mechanical properties, particularly the higher level strain-rate dependency, of the biomaterials used in the FE simulations of the human body, SHPB experiments were performed on the biomaterials to obtain dynamic mechanical responses at strain rates pertaining to human TIs. An overview of the SHPB setup at the Center for Advanced Vehicular Systems (CAVS), Mississippi State University (MSU) is presented in Figure 1.
Previous studies have shown that SHPB testing has three major flaws associated with it12-18. The first and most significant one is the material inertial effect, which shows up in the high strain rate mechanical response of a biomaterial specimen as an initial spike. In order to overcome this issue, previous research efforts suggested modifying the geometry of the specimen from cylindrical in shape to cuboidal or annular in shape. The resultant mechanical behaviors from such studies were different from each other because the geometry of the specimen affected the wave propagation, wave interactions, and the mechanical response. This type of modification to the specimen geometry has led to erroneous representations of the mechanical response (multiaxial and non-uniform stress state) of the biomaterial. The second major flaw was the inability to maintain dynamic force equilibrium during a test. Researchers overcame this issue by reducing the sample thickness-to-diameter ratio and/or freezing the tissue prior to testing. While reducing the sample thickness-to-diameter ratio addressed the issue of dynamic force equilibrium, freezing the tissue further complicated the testing procedure as it changed the material properties due to crystallization of water present in the tissue. A number of studies completely abandoned the SHPB to avoid the above mentioned flaws and used shock tubes to obtain the pressure-time response in various animal models (rats, pigs, etc.). However, these animal models do not give one-dimensional uniaxial stress-strain behaviors necessary for material models used in FE simulations. The third flaw was the failure of the SHPB to give one dimensional stress-strain results because of the specimen barreling due to the material softness and the amount of water content in the specimen.
Hence, the SHPB presents a viable testing apparatus to garner high strain rate data . For soft materials, however, the SHPB induces bulging that produces a three-dimensional stress state mainly from hydrostatic pressure, yet the one dimensional stress-strain data is desired. We show here how one can still use the SHPB to garner the one-dimensional uniaxial true stress-strain curve for material model calibration; however, the process involved in obtaining the uniaxial true stress-strain curve is complicated. This process includes both the multi-axial experimental data and FE simulation results, and it requires iterative recalibration of the material model constants. The one-dimensional implementation of the ISV material model in MATLAB, also known as material point simulator, requires one-dimensional experimental data for the calibration. So, the ISV material model was optimized using a systematic calibration process. Here, experimental data from SHPB tests was considered in the context of wave theory formulation and dynamic force equilibrium (MSU High Rate Software). In order to account for the viscoelastic dispersion of the polymeric SHPB, viscoelastic dispersion equations, as reported by Zhao et al. (2007), were implemented in MSU High Rate Software. The viscoelastic dispersion equations helped in ensuring dynamic force equilibrium while testing. The one-dimensional material point simulator was then adjusted in the context of a couple experiment-FE modeling methodology until the two processes were considered to be appropriately compatible, that is, the data from both were in good agreement. These data were used to adjust the ISV model material constants by comparing the MATLAB material response simulator&#8217;s (one-dimensional) mechanical response and the SHPB FE model&#8217;s (one-dimensional) specimen centerline stress. Here the FE model&#8217;s specimen stress component was along the wave loading direction. Then the three-dimensional behavior of the FE model specimen was calibrated by iteratively performing FE simulations and adjusting ISV constants so that volume-averaged loading direction stress correlated well with the experimental true stress-strain response. Thus, a process of iterative optimization between the experimental data, FE results, and one-dimensional ISV material model was conducted. Table 1 gives a summary of the variables of ISV material model (MSU TP Ver. 1.1)11.
The most important element to this methodology is obtaining the one-dimensional mechanical response of the biomaterial and its material parameters for the ISV material model, which circumvents the SHPB testing issues of the stress-state non-uniformity. It also separates out the initial nonlinear response of the biomaterial arising from inertial effects and renders a mechanical response that is intrinsic to the material. The coupled methodology also showed that a change in the specimen geometry completely changes the Boundary Value Problem (BVP) and the loading direction true stress-strain of the specimen. As such, the above mentioned methodology can be used with any material model (phenomenological or microstructural-based) for calibrating and then simulating high strain rate behaviors of human organs under injurious mechanical loads.

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			<td align="right">2</td>
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		<tr height="20">
			<td height="20" style="height:20px;">1/4&#34; x 25 ft polyethylene tubing</td>
			<td>Lowes</td>
			<td></td>
			<td align="right">2</td>
		</tr>
		<tr height="20">
			<td height="20" style="height:20px;">1-1/2&#34; Diameter Polycarbonate (PC) Rod</td>
			<td>McMaster-Carr</td>
			<td></td>
			<td align="right">2</td>
		</tr>
		<tr height="20">
			<td height="20" style="height:20px;">LTV-35 4-Way Valve Mead Fluid Dynamics</td>
			<td>Motion Industries</td>
			<td></td>
			<td align="right">2</td>
		</tr>
		<tr height="20">
			<td height="20" style="height:20px;">Pneumatic double action actuator</td>
			<td>Valtronic</td>
			<td></td>
			<td align="right">2</td>
		</tr>
		<tr height="20">
			<td height="20" style="height:20px;">Stainless Steel Ball Valve 1/2&#34;</td>
			<td>Valtronic</td>
			<td></td>
			<td align="right">2</td>
		</tr>
		<tr height="20">
			<td height="20" style="height:20px;">Buckeye pressure vessel</td>
			<td>Buckeye</td>
			<td></td>
			<td align="right">2</td>
		</tr>
		<tr height="60">
			<td height="60" style="height:60px;">SR-4 General Purpose FAE-25-35SX Strain Gages</td>
			<td style="width:141px;">Micro-Measurement Vishay Precision Group</td>
			<td style="width:53px;"></td>
			<td align="right">2</td>
		</tr>
		<tr height="60">
			<td height="60" style="height:60px;">M-M Signal Conditioning Amplifier 2310A</td>
			<td style="width:141px;">Micro-Measurement Vishay Precision Group</td>
			<td style="width:53px;"></td>
			<td align="right">1</td>
		</tr>
		<tr height="20">
			<td height="20" style="height:20px;">Laser ROLS-W optical sensor</td>
			<td>Monarch Instruments</td>
			<td></td>
			<td align="right">1</td>
		</tr>
	</tbody>
</table>

a coupled experiment-finite element modeling methodology for assessing high strain rate mechanical response of soft biomaterials

This study offers a combined experimental and finite element (FE) simulation approach for examining the mechanical behavior of soft biomaterials (e.g. brain, liver, tendon, fat, etc.) when exposed to high strain rates. This study utilized a Split-Hopkinson Pressure Bar (SHPB) to generate strain rates of 100-1,500 sec-1. The SHPB employed a striker bar consisting of a viscoelastic material (polycarbonate). A sample of the biomaterial was obtained shortly postmortem and prepared for SHPB testing. The specimen was interposed between the incident and transmitted bars, and the pneumatic components of the SHPB were activated to drive the striker bar toward the incident bar. The resulting impact generated a compressive stress wave (i.e. incident wave) that traveled through the incident bar. When the compressive stress wave reached the end of the incident bar, a portion continued forward through the sample and transmitted bar (i.e. transmitted wave) while another portion reversed through the incident bar as a tensile wave (i.e. reflected wave). These waves were measured using strain gages mounted on the incident and transmitted bars. The true stress-strain behavior of the sample was determined from equations based on wave propagation and dynamic force equilibrium. The experimental stress-strain response was three dimensional in nature because the specimen bulged. As such, the hydrostatic stress (first invariant) was used to generate the stress-strain response. In order to extract the uniaxial (one-dimensional) mechanical response of the tissue, an iterative coupled optimization was performed using experimental results and Finite Element Analysis (FEA), which contained an Internal State Variable (ISV) material model used for the tissue. The ISV material model used in the FE simulations of the experimental setup was iteratively calibrated (i.e. optimized) to the experimental data such that the experiment and FEA strain gage values and first invariant of stresses were in good agreement.

The effectiveness of the coupled methodology is exemplified in Figure 3. Here the SHPB experimental stress-strain response for the brain is at a lower stress state (with a peak stress of 0.32 MPa) in comparison to the stress state of the one-dimensional material point simulator (with a peak value of 0.74 MPa), which is akin to the FE sample center line (element) average. This is due to the nature of deformation that soft biomaterials exhibit. Because the strain rates are high, and the wave speed and stre...

Watch this Scientific Journal Video about A Coupled Experiment-finite Element Modeling Methodology for Assessing High Strain Rate Mechanical Response of Soft Biomaterials at JoVE.com

A Coupled Experiment-finite Element Modeling Methodology for Assessing High Strain Rate Mechanical Response of Soft Biomaterials

The current study prescribes a coupled experiment-finite element simulation methodology to obtain the uniaxial dynamic mechanical response of soft biomaterials (brain, liver, tendon, fat, etc.). The multiaxial experimental results that arose because of specimen bulging obtained from Split-Hopkinson Pressure Bar testing were rendered to a uniaxial true stress-strain behavior when simulated through iterative optimization of the finite element analysis of the biomaterial.

This study offers a combined experimental and finite element (FE) simulation approach for examining the mechanical behavior of soft biomaterials (e.g. ...

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