The authors thank Texas A&#38;M University Engineering and Experiment Station (TEES) for financial support of this research. Lindsay Nail gratefully acknowledges support from the Texas A&#38;M University Louis Stokes Alliance for Minority Participation (LSAMP) and the National Science Foundation (NSF) Graduate Research Fellowship Program (GRFP). Dawei Zhang thanks the Texas A&#38;M University Dissertation Fellowship.

1. Synthesizing PCL-DA Macromer
<ol>
	<li>Run the acrylation reaction.
	<ol>
		<li>Weigh 20 g of PCL-diol (Mn = ~10,000 g/mol) into a 250 ml&#160;round bottom flask equipped with a Teflon-covered magnetic stir bar.</li>
		<li>Dissolve the PCL-diol in DCM.
		<ol>
			<li>Add 120 mLlof DCM to the flask (concentration = 0.17 g/ml).</li>
			<li>Place a rubber septum loosely into the neck of the flask so as to avoid pressure build-up while also preventing evaporation of DCM.</li>
			<li>Stir solution for ~30 min&#160;at ~250 rpm to completely dissolve the polymer.</li>
		</ol>
		</li>
		<li>Add ~6.6 mg of 4-dimethylaminopyridine (DMAP) to the solution and dissolve with stirring.</li>
		<li>Place a rubber septum securely into the neck of the flask. Allow the solution to continue stirring.</li>
		<li>Through the rubber septum, gently purge the flask with N2 for ~3 min&#160;by using a positive N2 pressure needle inlet and an open needle as an outlet.</li>
		<li>Remove the N2 inlet and outlet.</li>
		<li>Add 0.56 ml&#160;(4.0 mmol) of triethylamine (Et3N) dropwise via a glass syringe equipped with a needle inserted through the rubber septum.</li>
		<li>Add 0.65 ml&#160;(8.0 mmol) of acryloyl chloride dropwise via a glass syringe equipped with a needle inserted through the rubber septum.</li>
		<li>Return the N2 inlet to the flask and allow the contents to stir under positive N2 pressure for ~30 min.</li>
		<li>Pre-heat an oil bath to 55 &#186;C.</li>
		<li>After the allotted ~30 min, remove the N2 inlet and replace the septum with a condenser.</li>
		<li>Submerge the flask into the pre-heated oil bath.</li>
		<li>Allow the contents of the flask to stir for 20 hr.</li>
		<li>After the allotted 20 hr, remove the flask from the oil bath and allow contents to cool to RT.</li>
		<li>Using a rotary evaporator, remove DCM solvent from the flask.</li>
	</ol>
	</li>
	<li>Purify the crude PCL-DA product.
	<ol>
		<li>To the flask, add ~135 ml&#160;of ethyl acetate and dissolve the crude PCL-DA.</li>
		<li>Gravity filter the solution through filter paper into a clean 250 ml&#160;round bottom flask. (Note: Solution may thicken on the filter paper, not readily passing through. If so, carefully apply mild heat with a heating gun.)</li>
		<li>Using a rotary evaporator, remove ethyl acetate solvent from the flask.</li>
		<li>To the flask, add ~140 ml&#160;of DCM and dissolve the crude PCL-DA.</li>
		<li>Transfer contents to a 500 ml&#160;separatory funnel.</li>
		<li>To the funnel, add 13.5 ml&#160;of 2 M potassium carbonate (K2CO3).</li>
		<li>Cap the funnel. Gently mix the two layers by inverting the funnel and swirling gently once or twice, taking care to release pressure via the stopcock. Repeat 3 times.</li>
		<li>Replace the cap with a layer of Parafilm and allow the mixture to separate O/N (~12 hr).</li>
		<li>Collect the bottom, organic layer into a 250 ml&#160;Erlenmeyer flask.</li>
		<li>Add ~5 g of anhydrous magnesium sulfate (MgSO4) to the flask and gently swirl.</li>
		<li>Gravity filter the mixture through qualitative filter paper and into a clean 250 ml&#160;round bottom flask.</li>
		<li>Using a rotary evaporator, remove DCM solvent from the flask.</li>
		<li>Dry under high vacuum to remove residual DCM. (Note: PCL-DA should be stored away from light.)</li>
		<li>Confirm acrylation with 1H NMR22,23.</li>
	</ol>
	</li>
</ol>
2. Preparing the SMP Scaffold (Figure 1)
<ol>
	<li>Prepare the fused salt template.
	<ol>
		<li>Use a 425 &#181;m sieve to obtain sodium chloride (NaCl) particles ~460 &#177; 70 &#181;m in diameter. (Note: Average particle size may be confirmed from scanning electron microscopy [SEM] images with ImageJ software .)14</li>
		<li>To a 3 ml&#160;glass vial (I.D. = 12.9 mm), add 1.8 g of the previously sieved NaCl.</li>
		<li>Slowly add, in four portions, 7.5 wt% (based on salt weight) DI water (0.146 g) to the vial. Mix with a metal spatula after the addition of each portion of water.</li>
		<li>Cap the vial, wrap in tissue and place vertically into a centrifuge tube. Centrifuge for 15 min&#160;at 3,220 x g.</li>
		<li>Remove the cap and let air-dry O/N (~12 hr).</li>
	</ol>
	</li>
	<li>In a new glass vial, prepare a &#8220;macromer solution&#8221; by combining 0.15 g of PCL-DA per ml&#160;of DCM. (Note: For one scaffold, ~ 1 ml&#160;of solution should be prepared.) Cap and mix the solution at high speed on a vortex mixer for ~1 min.</li>
	<li>In a new 3 ml&#160;glass vial, prepare a &#8220;photoinitiator solution&#8221; based on 10 wt% 2,2-dimethoxy-2-phenyl acetophenone (DMP) in 1-vinyl-2-pyrrolidinone (NVP). Combine 0.115 g of DMP in 1 ml&#160;of NVP. (Note: For one scaffold, ~180 &#181;l is required.) Cap, wrap in aluminum foil (to block light) and mix the solution at high speed on a vortex mixer for ~1 min. (Note: If prepared in advanced, the solution should be refrigerated and protected from light.)</li>
	<li>Wrap the vial containing the macromer solution (excluding the cap) with aluminum foil (to block light) and add via pipette 15 vol% (based on total volume of the macromer solution) of the photoinitiator solution. Cap and mix the solution at high speed on a vortex mixer for ~1 min.</li>
	<li>Wrap the vial containing the fused salt template (excluding the cap) with aluminum foil (to block light) and add via pipette the previously prepared macromer/photoinitiator solution (~0.6 ml&#160;or until the template is completely covered).</li>
	<li>Cap the vial, wrap in tissue and place vertically into a centrifuge tube. Centrifuge for 10 min&#160;at 1,260 x g to distribute the macromer solution throughout the template.</li>
	<li>Remove the aluminum foil, uncap the vial and expose to UV light (365 nm, 25 W) for 3 min. Air-dry O/N.</li>
	<li>Remove the &#8220;salt-containing scaffold&#8221; from the vial with tweezers after scoring and fracturing the top of the glass vial.</li>
	<li>In a 400 ml&#160;beaker, prepare ~200 ml&#160;of a water/ethanol solvent (1:1 vol:vol).</li>
	<li>Maintain the scaffold submerged in the water/ethanol solvent for 4 days with daily solvent changes.</li>
	<li>Remove the scaffold from the solvent and air-dry O/N.</li>
</ol>
3. Applying Polydopamine Coating to SMP Scaffold (Figure 1)
<ol>
	<li>In 400 ml&#160;beaker equipped with a Teflon-covered stir bar, prepare ~200 ml&#160;of a dopamine hydrochloride solution (2 mg/ml&#160;in 10 mM Tris buffer, pH = 8.5, 25 &#186;C). Stir at ~150 rpm.</li>
	<li>Place a disposable needle (length = 40 mm; gauge = 20) into the scaffold, ~half the distance through the scaffold. Wrap a wire around the needle hub.</li>
	<li>Submerge the scaffold (with the needle hub above the solution surface) into the stirring dopamine solution by anchoring the wire to the rim of the beaker.</li>
	<li>Degas the scaffold by placing a syringe into the needle hub and using it to pull air out of the scaffold. (Note: Degassing is complete when no more air can be removed and the solution has completely infiltrated the scaffold.)</li>
	<li>Maintain the scaffold submerged in the stirring dopamine solution for 16 hr.</li>
	<li>Remove the scaffold from the solution and remove the needle. Rinse with DI water and dry in a vacuum oven at RT for 24 hr.</li>
	<li>Place the scaffold in an 85 &#186;C oven for 1 hr.</li>
	<li>Allow the scaffold to cool to RT. The final cylindrical scaffold will be ~6 mm diameter x ~5 mm height.</li>
</ol>
4. Evaluating &#8220;Self-fitting&#8221; Behavior
<ol>
	<li>Prepare an &#8220;irregular CMF defect model&#8221; using a sheet of rigid plastic whose thickness is ~5 mm. Use a drill to create a void within the plastic sheet with an average diameter slightly less than ~6 mm, as demonstrated in Figure 2A.</li>
	<li>In a beaker, heat DI water (representing the clinical use of saline) to a temperature of ~60 &#186;C.</li>
	<li>Place the scaffold into the beaker of ~60 &#186;C water. Use tweezers to push the scaffold below the water surface, exposing all areas to the water. Continue for ~2 min or until the scaffold is noticeably softened (Figure 2B).</li>
	<li>Remove the scaffold from the beaker and immediately press (by hand) into the model defect.</li>
	<li>Allow to cool to RT (~5-10 min) (Figure 2C).</li>
	<li>Remove from the defect to observe the new, fixed temporary shape and the return of the relatively more rigid state (Figure 2D).</li>
</ol>
5. Testing Shape Memory Behavior
<ol>
	<li>Using a dynamic mechanical analyzer (DMA; e.g. a TA Instruments Q800 as used herein), run a strain-controlled cyclic-thermal mechanical compression test on a scaffold over two cycles (N) to determine shape fixity (Rf) and shape recovery (Rr) (Figure 3).
	<ol>
		<li>Equilibrate to 60 &#176;C (Thigh) for 5 min.</li>
		<li>Compress to a maximum strain (&#949;m = 50 %) at 50 %/min.</li>
		<li>Hold at &#949;m (5 min).</li>
		<li>Cool to 25 &#176;C (Tlow) and maintain for 10 min to fix the temporary shape.</li>
		<li>Remove the load.</li>
		<li>Measure the ultimate strain in the stress-free state (&#949;u).</li>
		<li>Reheat to 60 &#176;C (Thigh) and maintain for 10 min to recover the permanent shape.</li>
		<li>Measure the recovered strain (&#949;p).</li>
		<li>While still at 60 &#176;C (Thigh), start the 2nd cycle (N = 2) by compressing the scaffold to 50 % of the height recovered after the 1st cycle (N = 1).</li>
		<li>Repeat 5.1.3-5.1.8 for N = 2.</li>
		<li>Calculate Rf and Rr for N = 1 and 2 using the following equations: 
		Rf(N) = [&#949;u(N)/ &#949;m]Rr(N) = [&#949;m - &#949;p(N)]/[&#949;m- &#949;p(N-1)]</li>
	</ol>
	</li>
</ol>
6. Visualizing Pore Size and Pore Interconnectivity
<ol>
	<li>Using scanning electron microscopy (SEM; e.g. a FEI Quanta SEM as used herein), observe the pore size and pore interconnectivity.
	<ol>
		<li>Using tweezers to hold the SMP scaffold, submerge in liquid N2 for 1 min.</li>
		<li>Remove from liquid N2 and fracture along the middle of the scaffold with a clean razor blade.</li>
		<li>Using carbon tape, affix one of the SMP scaffold halves onto the sample stage with the fractured surface facing up.</li>
		<li>Sputter coat with Au-Pt (~4 nm).</li>
		<li>Capture the SEM image at a recommended accelerating voltage of 10-15 kV ( Figure 4A).</li>
	</ol>
	</li>
</ol>
7. Testing of in vitro Bioactivity
<ol>
	<li>Into a 50 ml&#160;centrifuge tube, add ~30 ml&#160;of 1X SBF24.</li>
	<li>Obtain a scaffold in its original, cylindrically-molded permanent shape. Cut the scaffold in half (across the circular edge) using a clean blade.</li>
	<li>Place an individual scaffold half into the prepared centrifuge tube and cap.</li>
	<li>Maintain the tube at 37 &#176;C in a water bath under static conditions with no SBF changes.</li>
	<li>After 14 days, remove the scaffold from the SBF and air dry for 24 hr.</li>
	<li>Using carbon tape, affix the scaffold onto the sample stage with the fractured surface facing up.</li>
	<li>Sputter coat with Au-Pt (~4 nm).</li>
	<li>Capture the SEM image at a recommended accelerating voltage of 10-15 kV (Figure 4B).</li>
</ol>

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The authors have nothing to disclose.

This protocol describes the preparation of a polydopamine-coated, PCL-based scaffold whose self-fitting behavior, as well as osteoinductivity and bioactivity, makes it of interest in the treatment of irregular CMF bone defects. Aspects of the protocol may be altered to change various scaffold features.
The protocol begins with acrylation of a PCL-diol to permit UV cure. In the reported example, the PCL-diol Mn is ~10,000 g/mol. However, by appropriately adjusting amount of acryloyl ...

Currently considered the gold standard of cranio-maxillofacial (CMF) bone defect treatments, transplantation of harvested autologous grafts is hindered by complex grafting procedures, donor site morbidity and limited availability1. A particular difficulty is shaping and fixing the rigid autograft tightly into the defect in order to obtain osseointegration and to prevent graft resorption. Tissue engineering has been investigated as an alternative strategy to autografting and synthetic bone substitutes (e.g. bone cement)2,3. Critical to the success of a tissue engineering approach is a scaffold with a specific set of properties. First, in order to achieve osseointegration, the scaffold must form close contact with adjacent bone tissue4. The scaffold should also be osteoconductive, permitting cell migration, nutrient diffusion and neotissue deposition4,5. This behavior is generally achieved with biodegradable scaffolds exhibiting a highly interconnected pore morphology. Lastly, the scaffold should be bioactive so as to promote integration and bonding with surrounding bone tissue5.
Here, we present a protocol to prepare a tissue engineering scaffold with these properties. Importantly, this scaffold exhibits the ability to &#8220;self-fit&#8221; into irregular CMF defects due to its shape memory behavior6. Thermoresponsive shape memory polymers (SMPs) are known to undergo shape change upon exposure to heat7,8. SMPs are comprised of &#8220;netpoints&#8221; (i.e. chemical or physical crosslinks) which determine the permanent shape and &#8220;switching segments&#8221; which maintain the temporary shape and recover the permanent shape. The switching segments exhibit a thermal transition temperature (Ttrans) corresponding to either the glass transition (Tg) or melt transition (Tm) of the polymer. As a result, SMPs may be sequentially deformed into a temporary shape at T &#62; Ttrans, fixed in the temporary shape at T &#60; Ttrans, and recovered to the permanent shape at T &#62; Ttrans. Thus, an SMP scaffold could achieve &#8220;self-fitting&#8221; within a CMF defect as follows6. After exposure to warm saline (T &#62; Ttrans), an SMP scaffold would become malleable, permitting a generically prepared cylindrical scaffold to be hand-pressed into an irregular defect, with shape recovery promoting expansion of the scaffold to the defect boundary. Upon cooling (T &#60; Ttrans), the scaffold would return to its relatively more rigid state, with shape fixity maintaining its new temporary shape within the defect. In this protocol, an SMP scaffold is prepared from polycaprolactone (PCL), a biodegradable polymer studied extensively for tissue regeneration and other biomedical applications9-11. For shape memory, the Tm of PCL serves as the Ttrans and varies between 43 and 60 &#186;C, depending on the molecular weight of the PCL12. In this protocol, the Ttrans (i.e. Tm) of the scaffold is 56.6&#177;0.3 &#186;C6.
In order to achieve osteoconductivity, a protocol was developed to make PCL-based SMP scaffolds with highly interconnected pores based on a solvent-casting particulate-leaching (SCPL) method6,13,14. Polycaprolactone diacrylate (PCL-DA) (Mn = ~10,000 g/mol) was utilized to permit rapid, photochemical crosslinking and was dissolved in dichloromethane (DCM) to allow solvent-casting over the salt template. Following photochemical cure and solvent evaporation, the salt template was removed by leaching into water. The average salt size regulates scaffold pore size. Importantly, the salt template was fused with water prior to solvent-casting to achieve pore interconnectivity.
Bioactivity was imparted to the SMP scaffold by the in situ formation of a polydopamine coating onto pore walls6. Bioactivity is often introduced into scaffolds by the inclusion of glass or glass-ceramic fillers15. However, these may give rise to unwanted brittle mechanical properties. Dopamine has been shown to form an adherent, thin polydopamine layer on a variety of substrates16-19. In this protocol, the SMP scaffold was subjected to a slightly basic solution (pH = 8.5) of dopamine to form a nanothick coating of polydopamine on all pore wall surfaces6. In addition to enhancing surface hydrophilicity for improved cell adhesion and spreading, polydopamine has been shown to be bioactive in terms of formation of hydroxyapatite (HAp) upon exposure to simulated body fluid (SBF)18,20,21. In a last step, the coated scaffold is exposed to heat treatment at 85 &#186;C (T &#62; Ttrans) which leads to scaffold densification. Heat treatment was previously noted to be essential for scaffold shape memory behavior, perhaps due to PCL crystalline domains reorganizing to closer proximity14.
We additionally describe the methods to characterize the self-fitting behavior within an irregular model defect, shape memory behavior in terms strain-controlled cyclic-thermal mechanical compression tests (i.e. shape recovery and shape fixity), pore morphology, and in vitro bioactivity. Strategies to tailor scaffold properties are also presented.

<table border="0" cellpadding="0" cellspacing="0">
	<tbody>
		<tr>
			<td>Polycaprolactone-diol (Mn ~ 10,000 g/mol)</td>
			<td>Sigma-Aldrich</td>
			<td>440752</td>
			<td></td>
		</tr>
		<tr>
			<td>Dichloromethane (DCM)</td>
			<td>Sigma-Aldrich</td>
			<td>D65100</td>
			<td>Dried over 4A molecular sieves</td>
		</tr>
		<tr>
			<td>4-dimethylaminopyridine (DMAP)</td>
			<td>Sigma-Aldrich</td>
			<td>D5640</td>
			<td></td>
		</tr>
		<tr>
			<td>Triethylamine (Et3N)</td>
			<td>Sigma-Aldrich</td>
			<td>T0886</td>
			<td></td>
		</tr>
		<tr>
			<td>Acryloyl chloride</td>
			<td>Sigma-Aldrich</td>
			<td>A24109</td>
			<td></td>
		</tr>
		<tr>
			<td>Ethyl Acetate</td>
			<td>Sigma-Aldrich</td>
			<td>319902</td>
			<td></td>
		</tr>
		<tr>
			<td>Potassium Carbonate (K2CO3)</td>
			<td>Sigma-Aldrich</td>
			<td>209619</td>
			<td></td>
		</tr>
		<tr>
			<td>Anhydrous magnesium sulfate (MgSO4)</td>
			<td>Fisher</td>
			<td>M65</td>
			<td></td>
		</tr>
		<tr>
			<td>Sodium chloride (NaCl)</td>
			<td>Sigma-Aldrich</td>
			<td>S9888</td>
			<td></td>
		</tr>
		<tr>
			<td>2,2-dimethoxy-2-phenyl acetophenone (DMP)</td>
			<td>Sigma-Aldrich</td>
			<td>196118</td>
			<td></td>
		</tr>
		<tr>
			<td>1-vinyl-2-pyrrolidinone (NVP)</td>
			<td>Sigma-Aldrich</td>
			<td>V3409</td>
			<td></td>
		</tr>
		<tr>
			<td>Ethanol</td>
			<td>Sigma-Aldrich</td>
			<td>459844</td>
			<td></td>
		</tr>
		<tr>
			<td>Dopamine Hydrochloride</td>
			<td>Sigma-Aldrich</td>
			<td>H8502</td>
			<td></td>
		</tr>
		<tr>
			<td>Tris buffer (2mol/L)</td>
			<td>Fisher</td>
			<td>BP1759</td>
			<td>Used at 10 mM concentration, pH = 8.5</td>
		</tr>
		<tr>
			<td>Sieve</td>
			<td>VWR</td>
			<td>47729-972</td>
			<td></td>
		</tr>
		<tr>
			<td>UV-Transilluminator (365 nm, 25 W)</td>
			<td>UVP</td>
			<td>95-0426-02</td>
			<td></td>
		</tr>
		<tr>
			<td>Centrifuge</td>
			<td>Eppendorf</td>
			<td>5810 R</td>
			<td></td>
		</tr>
		<tr>
			<td>Dynamic Mechanical Analyzer (DMA)</td>
			<td>TA Instruments</td>
			<td>Q800</td>
			<td></td>
		</tr>
		<tr>
			<td>High Resolution Sputter Coater</td>
			<td>Cressington</td>
			<td>208HR</td>
			<td></td>
		</tr>
		<tr>
			<td>Scanning Electron Microscope (SEM)</td>
			<td>FEI</td>
			<td>Quanta 600</td>
			<td></td>
		</tr>
	</tbody>
</table>

fabrication of a bioactive, pcl-based "self-fitting" shape memory polymer scaffold

Tissue engineering has been explored as an alternative strategy for the treatment of critical-sized cranio-maxillofacial (CMF) bone defects. Essential to the success of this approach is a scaffold that is able to conformally fit within an irregular defect while also having the requisite biodegradability, pore interconnectivity and bioactivity. By nature of their shape recovery and fixity properties, shape memory polymer (SMP) scaffolds could achieve defect &#8220;self-fitting.&#8221; In this way, following exposure to warm saline (~60 &#186;C), the SMP scaffold would become malleable, permitting it to be hand-pressed into an irregular defect. Subsequent cooling (~37 &#186;C) would return the scaffold to its relatively rigid state within the defect. To meet these requirements, this protocol describes the preparation of SMP scaffolds prepared via the photochemical cure of biodegradable polycaprolactone diacrylate (PCL-DA) using a solvent-casting particulate-leaching (SCPL) method. A fused salt template is utilized to achieve pore interconnectivity. To realize bioactivity, a polydopamine coating is applied to the surface of the scaffold pore walls. Characterization of self-fitting and shape memory behaviors, pore interconnectivity and in vitro bioactivity are also described.

The resulting PCL-based SMP scaffold is capable of self-fitting into a model CMF defect (Figure 2). After brief exposure to warm saline (~60 &#176;C), the cylindrical scaffold softens allowing the scaffold to be manually pressed into and expand within the model defect. After cooling to RT, the scaffold is fixed into its new temporary shape which is retained upon removal from the defect.
The shape memory behavior of an SMP scaffold is quantified by strain-controlled cyclic-ther...

Watch this Scientific Journal Video about Fabrication of a Bioactive, PCL-based "Self-fitting" Shape Memory Polymer Scaffold at JoVE.com

Fabrication of a Bioactive, PCL-based &quot;Self-fitting&quot; Shape Memory Polymer Scaffold

Scaffolds capable of fitting within cranio-maxillofacial (CMF) bone defects while exhibiting osteoconductivity and bioactivity are of interest. This protocol describes the preparation of a shape memory scaffold based on polycaprolactone diacrylate (PCL-DA) using a solvent-casting particulate-leaching (SCPL) method employing a fused salt template and application of a bioactive polydopamine coating.

Fabrication of a Bioactive, PCL-based "Self-fitting" Shape Memory Polymer Scaffold

Tissue engineering has been explored as an alternative strategy for the treatment of critical-sized cranio-maxillofacial (CMF) bone defects. Essential to ...

fabrication-bioactive-pcl-based-self-fitting-shape-memory-polymer

Research

JoVE Journal

Bioengineering

12.5K 조회수.  Texas A&M University. Scaffolds capable of fitting within cranio-maxillofacial (CMF) bone defects while exhibiting osteoconductivity and bioactivity are of interest. This protocol describes the preparation of a shape memory scaffold based on polycaprolactone diacrylate (PCL-DA) using a solvent-casting particulate-leaching (SCPL) method employing a fused salt template and application of a bioactive polydopamine coating.

비디오: Fabrication of a Bioactive, PCL-based "Self-fitting" Shape Memory Polymer Scaffold

Shape Memory Polymers for Active Cell Culture

Shape memory polymers (SMPs) are a class of "smart" materials that have the ability to change from a fixed, temporary shape to a pre-determined permanent shape upon the application of a stimulus such as heat1-5. In a typical shape memory cycle, the SMP is first deformed at an elevated temperature that is higher than its transition temperature, Ttrans [either the melting temperature (Tm) or the glass transition temperature (Tg)]. The deformation is elastic in nature and mainly leads to a reduction in conformational entropy of the constituent network chains (following the rubber elasticity theory). The deformed SMP is then cooled to a temperature below its Ttrans while maintaining the external strain or stress constant. During cooling, the material transitions to a more rigid state (semi-crystalline or glassy), which kinetically traps or "freezes" the material in this low-entropy state leading to macroscopic shape fixing. Shape recovery is triggered by continuously heating the material through Ttrans under a stress-free (unconstrained) condition. By allowing the network chains (with regained mobility) to relax to their thermodynamically favored, maximal-entropy state, the material changes from the temporary shape to the permanent shape.
Cells are capable of surveying the mechanical properties of their surrounding environment6. The mechanisms through which mechanical interactions between cells and their physical environment control cell behavior are areas of active research. Substrates of defined topography have emerged as powerful tools in the investigation of these mechanisms. Mesoscale, microscale, and nanoscale patterns of substrate topography have been shown to direct cell alignment, cell adhesion, and cell traction forces7-14. These findings have underscored the potential for substrate topography to control and assay the mechanical interactions between cells and their physical environment during cell culture, but the substrates used to date have generally been passive and could not be programmed to change significantly during culture. This physical stasis has limited the potential of topographic substrates to control cells in culture.
Here, active cell culture (ACC) SMP substrates are introduced that employ surface shape memory to provide programmed control of substrate topography and deformation. These substrates demonstrate the ability to transition from a temporary grooved topography to a second, nearly flat memorized topography. This change in topography can be used to control cell behavior under standard cell culture conditions.

A method for developing cell culture substrates with the ability to change topography during culture is described. The method makes use of smart materials known as shape memory polymers that have the ability to memorize a permanent shape. This concept is adaptable to a wide range of materials and applications.

Shape memory polymers (SMPs) are a class of "smart" materials that have the ability to change from a fixed, temporary shape to a pre-determined permanent ...

A Chitosan Based, Laser Activated Thin Film Surgical Adhesive, 'SurgiLux': Preparation and Demonstration

Sutures are a 4,000 year old technology that remain the 'gold-standard' for wound closure by virtue of their repair strength (~100 KPa). However, sutures can act as a nidus for infection and in many procedures are unable to effect wound repair or interfere with functional tissue regeneration.1 Surgical glues and adhesives, such as those based on fibrin and cyanoacrylates, have been developed as alternatives to sutures for the repair of such wounds. However, current commercial adhesives also have significant disadvantages, ranging from viral and prion transfer and a lack of repair strength as with the fibrin glues, to tissue toxicity and a lack of biocompatibility for the cyanoacrylate based adhesives. Furthermore, currently available surgical adhesives tend to be gel-based and can have extended curing times which limit their application.2 Similarly, the use of UV lasers to facilitate cross-linking mechanisms in protein-based or albumin 'solders' can lead to DNA damage while laser tissue welding (LTW) predisposes thermal damage to tissues.3 Despite their disadvantages, adhesives and LTW have captured approximately 30% of the wound closure market reported to be in excess of US $5 billion per annum, a significant testament to the need for sutureless technology.4
In the pursuit of sutureless technology we have utilized chitosan as a biomaterial for the development of a flexible, thin film, laser-activated surgical adhesive termed 'SurgiLux'. This novel bioadhesive uses a unique combination of biomaterials and photonics that are FDA approved and successfully used in a variety of biomedical applications and products. SurgiLux overcomes all the disadvantages associated with sutures and current surgical adhesives (see Table 1).
In this presentation we report the relatively simple protocol for the fabrication of SurgiLux and demonstrate its laser activation and tissue weld strength. SurgiLux films adhere to collagenous tissue without chemical modification such as cross-linking and through irradiation using a comparatively low-powered (120 mW) infrared laser instead of UV light. Chitosan films have a natural but weak adhesive attraction to collagen (~3 KPa), laser activation of the chitosan based SurgiLux films emphasizes the strength of this adhesion through polymer chain interactions as a consequence of transient thermal expansion.5 Without this 'activation' process, SurgiLux films are readily removed.6-9 SurgiLux has been tested both in vitro and in vivo on a variety of tissues including nerve, intestine, dura mater and cornea. In all cases it demonstrated good biocompatibility and negligible thermal damage as a consequence of irradiation.6-10

A Chitosan Based, Laser Activated Thin Film Surgical Adhesive, &#039;SurgiLux&#039;: Preparation and Demonstration

The fabrication of a novel, flexible thin film surgical adhesive from FDA approved ingredients, chitosan and indocyanine green is described. Bonding of this adhesive to collagenous tissue through a simple activation process with a low-powered infra-red laser is demonstrated.

Sutures are a 4,000 year old technology that remain the 'gold-standard' for wound closure by virtue of their repair strength (~100 KPa). However, sutures ...

Environmentally-controlled Microtensile Testing of Mechanically-adaptive Polymer Nanocomposites for ex vivo Characterization

Implantable microdevices are gaining significant attention for several biomedical applications1-4. Such devices have been made from a range of materials, each offering its own advantages and shortcomings5,6. Most prominently, due to the microscale device dimensions, a high modulus is required to facilitate implantation into living tissue. Conversely, the stiffness of the device should match the surrounding tissue to minimize induced local strain7-9. Therefore, we recently developed a new class of bio-inspired materials to meet these requirements by responding to environmental stimuli with a change in mechanical properties10-14. Specifically, our poly(vinyl acetate)-based nanocomposite (PVAc-NC) displays a reduction in stiffness when exposed to water and elevated temperatures (e.g. body temperature). Unfortunately, few methods exist to quantify the stiffness of materials in vivo15, and mechanical testing outside of the physiological environment often requires large samples inappropriate for implantation. Further, stimuli-responsive materials may quickly recover their initial stiffness after explantation. Therefore, we have developed a method by which the mechanical properties of implanted microsamples can be measured ex vivo, with simulated physiological conditions maintained using moisture and temperature control13,16,17. 
To this end, a custom microtensile tester was designed to accommodate microscale samples13,17 with widely-varying Young's moduli (range of 10 MPa to 5 GPa). As our interests are in the application of PVAc-NC as a biologically-adaptable neural probe substrate, a tool capable of mechanical characterization of samples at the microscale was necessary. This tool was adapted to provide humidity and temperature control, which minimized sample drying and cooling17. As a result, the mechanical characteristics of the explanted sample closely reflect those of the sample just prior to explantation.
The overall goal of this method is to quantitatively assess the in vivo mechanical properties, specifically the Young's modulus, of stimuli-responsive, mechanically-adaptive polymer-based materials. This is accomplished by first establishing the environmental conditions that will minimize a change in sample mechanical properties after explantation without contributing to a reduction in stiffness independent of that resulting from implantation. Samples are then prepared for implantation, handling, and testing (Figure 1A). Each sample is implanted into the cerebral cortex of rats, which is represented here as an explanted rat brain, for a specified duration (Figure 1B). At this point, the sample is explanted and immediately loaded into the microtensile tester, and then subjected to tensile testing (Figure 1C). Subsequent data analysis provides insight into the mechanical behavior of these innovative materials in the environment of the cerebral cortex.

Environmentally-controlled Microtensile Testing of Mechanically-adaptive Polymer Nanocomposites for ex vivo Characterization

A method is discussed by which the in vivo mechanical behavior of stimuli-responsive materials is monitored as a function of time. Samples are tested ex vivo using a microtensile tester with environmental controls to simulate the physiological environment. This work further promotes understanding the in vivo behavior of our material.

Implantable microdevices are gaining significant attention for several biomedical applications1-4. Such devices have been made from a range of materials, ...

Microfluidic Fabrication of Polymeric and Biohybrid Fibers with Predesigned Size and Shape

A &#8220;sheath&#8221; fluid passing through a microfluidic channel at low Reynolds number can be directed around another &#8220;core&#8221; stream and used to dictate the shape as well as the diameter of a core stream.&#160;Grooves in the top and bottom of a microfluidic channel were designed to direct the sheath fluid and shape the core fluid.&#160;By matching the viscosity and hydrophilicity of the sheath and core fluids, the interfacial effects are minimized and complex fluid shapes can be formed.&#160;Controlling the relative flow rates of the sheath and core fluids determines the cross-sectional area of the core fluid.&#160;Fibers have been produced with sizes ranging from 300 nm to ~1 mm, and fiber cross-sections can be round, flat, square, or complex as in the case with double anchor fibers. Polymerization of the core fluid downstream from the shaping region solidifies the fibers.&#160;Photoinitiated click chemistries are well suited for rapid polymerization of the core fluid by irradiation with ultraviolet light.&#160;Fibers with a wide variety of shapes have been produced from a list of polymers including liquid crystals, poly(methylmethacrylate), thiol-ene and thiol-yne resins, polyethylene glycol, and hydrogel derivatives. Minimal shear during the shaping process and mild polymerization conditions also makes the fabrication process well suited for encapsulation of cells and other biological components.

Two adjacent fluids passing through a grooved microfluidic channel can be directed to form a sheath around a prepolymer core; thereby&#160;determining both shape and cross-section. Photoinitiated polymerization, such as thiol click chemistry, is well suited for rapidly solidifying the core fluid into a microfiber with predetermined size and shape.

A &#8220;sheath&#8221; fluid passing through a microfluidic channel at low Reynolds number can be directed around another &#8220;core&#8221; stream and ...

Manufacturing Of Robust Natural Fiber Preforms Utilizing Bacterial Cellulose as Binder

A novel method of manufacturing rigid and robust natural fiber preforms is presented here. This method is based on a papermaking process, whereby loose and short sisal fibers are dispersed into a water suspension containing bacterial cellulose. The fiber and nanocellulose suspension is then filtered (using vacuum or gravity) and the wet filter cake pressed to squeeze out any excess water, followed by a drying step. This will result in the hornification of the bacterial cellulose network, holding the loose natural fibers together.
Our method is specially suited for the manufacturing of rigid and robust preforms of hydrophilic fibers. The porous and hydrophilic nature of such fibers results in significant water uptake, drawing in the bacterial cellulose dispersed in the suspension. The bacterial cellulose will then be filtered against the surface of these fibers, forming a bacterial cellulose coating. When the loose fiber-bacterial cellulose suspension is filtered and dried, the adjacent bacterial cellulose forms a network and hornified to hold the otherwise loose fibers together.
The introduction of bacterial cellulose into the preform resulted in a significant increase of the mechanical properties of the fiber preforms. This can be attributed to the high stiffness and strength of the bacterial cellulose network. With this preform, renewable high performance hierarchical composites can also be manufactured by using conventional composite production methods, such as resin film infusion (RFI) or resin transfer molding (RTM). Here, we also describe the manufacturing of renewable hierarchical composites using double bag vacuum assisted resin infusion.

We present a novel method of manufacturing rigid and robust short natural fiber preforms using a papermaking process. Bacterial cellulose acts simultaneously as the binder for the loose fibers and provides rigidity to the fiber preforms. These preforms can be infused with a resin to produce truly green hierarchical composites.

A novel method of manufacturing rigid and robust natural fiber preforms is presented here. This method is based on a papermaking process, whereby loose ...

Fabrication of Mechanically Tunable and Bioactive Metal Scaffolds for Biomedical Applications

Biometal systems have been widely used for biomedical applications, in particular, as load-bearing materials. However, major challenges are high stiffness and low bioactivity of metals. In this study, we have developed a new method towards fabricating a new type of bioactive and mechanically reliable porous metal scaffolds-densified porous Ti scaffolds. The method consists of two fabrication processes, 1) the fabrication of porous Ti scaffolds by dynamic freeze casting, and 2) coating and densification of the porous scaffolds. The dynamic freeze casting method to fabricate porous Ti scaffolds allowed the densification of porous scaffolds by minimizing the chemical contamination and structural defects. The densification process is distinctive for three reasons. First, the densification process is simple, because it requires a control of only one parameter (degree of densification). Second, it is effective, as it achieves mechanical enhancement and sustainable release of biomolecules from porous scaffolds. Third, it has broad applications, as it is also applicable to the fabrication of functionally graded porous scaffolds by spatially varied strain during densification.

Bioactive and mechanically reliable metal scaffolds have been fabricated through a method which consists of two processes, dynamic freeze casting for the fabrication of porous Ti, and coating and densification of the Ti scaffolds. The densification process is simple, effective and applicable to the fabrication of functionally graded scaffolds.

Biometal systems have been widely used for biomedical applications, in particular, as load-bearing materials. However, major challenges are high stiffness ...

Fabrication of Inverted Colloidal Crystal Poly(ethylene glycol) Scaffold: A Three-dimensional Cell Culture Platform for Liver Tissue Engineering

The ability to maintain hepatocyte function in vitro, for the purpose of testing xenobiotics&#39; cytotoxicity, studying virus infection and developing drugs targeted at the liver, requires a platform in which cells receive proper biochemical and mechanical cues. Recent liver tissue engineering systems have employed three-dimensional (3D) scaffolds composed of synthetic or natural hydrogels, given their high water retention and their ability to provide the mechanical stimuli needed by the cells. There has been growing interest in the inverted colloidal crystal (ICC) scaffold, a recent development, which allows high spatial organization, homotypic and heterotypic cell interaction, as well as cell-extracellular matrix (ECM) interaction. Herein, we describe a protocol to fabricate the ICC scaffold using poly (ethylene glycol) diacrylate (PEGDA) and the particle leaching method. Briefly, a lattice is made from microsphere particles, after which a pre-polymer solution is added, properly polymerized, and the particles are then removed, or leached, using an organic solvent (e.g., tetrahydrofuran). The dissolution of the lattice results in a highly porous scaffold with controlled pore sizes and interconnectivities that allow media to reach cells more easily. This unique structure allows high surface area for the cells to adhere to as well as easy communication between pores, and the ability to coat the PEGDA ICC scaffold with proteins also shows a marked effect on cell performance. We analyze the morphology of the scaffold as well as the hepatocarcinoma cell (Huh-7.5) behavior in terms of viability and function to explore the effect of ICC structure and ECM coatings. Overall, this paper provides a detailed protocol of an emerging scaffold that has wide applications in tissue engineering, especially liver tissue engineering.

Fabrication of Inverted Colloidal Crystal Polyethylene glycol Scaffold: A Three-dimensional Cell Culture Platform for Liver Tissue Engineering

This manuscript presents a detailed protocol for the fabrication of an emerging three-dimensional hepatocyte culture platform, the inverted colloidal crystal scaffold, and the concomitant techniques to assess hepatocyte behavior. The size-controllable pores, interconnectivity and ability to conjugate extracellular matrix proteins to the poly(ethylene glycol) (PEG) scaffold enhance Huh-7.5 cell performance.

The ability to maintain hepatocyte function in vitro, for the purpose of testing xenobiotics&#39; cytotoxicity, studying virus infection and developing ...

Fabrication and Characterization of Optical Tissue Phantoms Containing Macrostructure

The rapid development of new optical imaging techniques is dependent on the availability of low-cost, customizable, and easily reproducible standards. By replicating the imaging environment, costly animal experiments to validate a technique may be circumvented. Predicting and optimizing the performance of in vivo and ex vivo imaging techniques requires testing on samples that are optically similar to tissues of interest. Tissue-mimicking optical phantoms provide a standard for evaluation, characterization, or calibration of an optical system. Homogenous polymer optical tissue phantoms are widely used to mimic the optical properties of a specific tissue type within a narrow spectral range. Layered tissues, such as the epidermis and dermis, can be mimicked by simply stacking these homogenous slab phantoms. However, many in vivo imaging techniques are applied to more spatially complex tissue where three dimensional structures, such as blood vessels, airways, or tissue defects, can affect the performance of the imaging system. 
This protocol describes the fabrication of a tissue-mimicking phantom that incorporates three-dimensional structural complexity using material with optical properties of tissue. Look-up tables provide India ink and titanium dioxide recipes for optical absorption and scattering targets. Methods to characterize and tune the material optical properties are described. The phantom fabrication detailed in this article has an internal branching mock airway void; however, the technique can be broadly applied to other tissue or organ structures.

Optical tissue phantoms are essential tools for calibration and characterization of optical imaging systems and validation of theoretical models. This article details a method for phantom fabrication that includes replication of tissue optical properties and three-dimensional tissue structure.

The rapid development of new optical imaging techniques is dependent on the availability of low-cost, customizable, and easily reproducible standards. By ...

Microdissection of Primary Renal Tissue Segments and Incorporation with Novel Scaffold-free Construct Technology

Kidney transplantation is now a mainstream therapy for end-stage renal disease. However, with approximately 96,000 people on the waiting list and only one-fourth of these patients achieving transplantation, there is a dire need for alternatives for those with failing organs. In order to decrease the harmful consequences of dialysis along with the overall healthcare costs it incurs, active investigation is ongoing in search of alternative solutions to organ transplantation. Implantable tissue-engineered renal cellular constructs are one such feasible approach to replacing lost renal functionality. Here, described for the first time, is the microdissection of murine kidneys for isolation of living corticomedullary renal segments. These segments are capable of rapid incorporation within scaffold-free endothelial-fibroblast constructs which may enable rapid connection with host vasculature once implanted. Adult mouse kidneys were procured from living donors, followed by stereoscope microdissection to obtain renal segments 200 - 300 &#181;m in diameter. Multiple renal constructs were fabricated using primary renal segments harvested from only one kidney. This method demonstrates a procedure which could salvage functional renal tissue from organs that would otherwise be discarded.

Tissue-engineered renal constructs provide a solution for the organ shortage and deleterious effects of dialysis. Here, we describe a protocol to micro dissect murine kidneys for isolation of cortico-medullary segments. These segments are implanted into scaffold-free cellular constructs, forming renal organoids.

Kidney transplantation is now a mainstream therapy for end-stage renal disease. However, with approximately 96,000 people on the waiting list and only ...

Scalable Fabrication of Stretchable, Dual Channel, Microfluidic Organ Chips

A significant number of lead compounds fail in the pharmaceutical pipeline because animal studies often fail to predict clinical responses in human patients. Human Organ-on-a-Chip (Organ Chip) microfluidic cell culture devices, which provide an experimental in vitro platform to assess efficacy, toxicity, and pharmacokinetic (PK) profiles in humans, may be better predictors of therapeutic efficacy and safety in the clinic compared to animal studies. These devices may be used to model the function of virtually any organ type and can be fluidically linked through common endothelium-lined microchannels to perform in vitro studies on human organ-level and whole body-level physiology without having to conduct experiments on people. These Organ Chips consist of two perfused microfluidic channels separated by a permeable elastomeric membrane with organ-specific parenchymal cells on one side and microvascular endothelium on the other, which can be cyclically stretched to provide organ-specific mechanical cues (e.g., breathing motions in lung). This protocol details the fabrication of flexible, dual channel, Organ Chips through casting of parts using 3D printed molds, enabling combination of multiple casting and post-processing steps. Porous poly (dimethyl siloxane) (PDMS) membranes are cast with micrometer sized through-holes using silicon pillar arrays under compression. Fabrication and assembly of Organ Chips involves equipment and steps that can be implemented outside of a traditional cleanroom. This protocol provides researchers with access to Organ Chip technology for in vitro organ- and body-level studies in drug discovery, safety and efficacy testing, as well as mechanistic studies of fundamental biological processes.

Here, we present a protocol that describes the fabrication of stretchable, dual channel, organ chip microfluidic cell culture devices for recapitulating organ-level functionality in vitro.

A significant number of lead compounds fail in the pharmaceutical pipeline because animal studies often fail to predict clinical responses in human ...