目前的项目是由堪萨斯大学概念基金的证明资助的。我们要感谢 KUMC (堪萨斯城, 堪萨斯州) 的护理人员帮助我们获得人类脐带, 和大卫乔根森协助耳蜗文化。

所有程序, 包括动物安乐死, 都是根据批准的机构动物保护和使用委员会 (IACUC) 议定书 (针灸学 #2014-2234) 在堪萨斯大学医学中心 (KUMC) 进行的。
注: HWJCs 是从人类脐带中分离出来的, 这些患者提供了知情同意, 并按照堪萨斯大学人类科目委员会 (KU-IRB #15402) 批准的协议使用了标本。
1. 颞骨收获与耳蜗隔离
<ol>
	<li>斩首一个适当的安乐死, 15 周大的老鼠通过切割头骨和第一个颈椎之间的一条锋利的外科剪刀, 然后喷雾下来的头与70% 乙醇消毒 (图 1A)。</li>
	<li>用同样锋利的外科剪刀 (图 1B-c, 虚线) 在 mid-sagittal 平面上对头骨。</li>
	<li>使用一对钳 (图 1D, 箭头) 从颅骨中取出脑组织。</li>
	<li>通过存在的听觉和前庭神经根 (图 1E-F, 箭头) 的内部头骨和耳管从外部 (图 1F, 箭头) 来识别颞骨。仔细地切开头骨从头骨的剩下的人分离世俗骨头 (图 1G)。 注意: 在某些情况下, 它可能可以微妙地撬开颅骨周围的骨骼, 而不需要切割。</li>
	<li>将细钳插入耳道的开口大约5毫米 (图 1H), 所以这些小贴士不会冒险刺穿耳蜗。轻轻地打开大的骨骼 (图 1H, 顶部红色阴影区域), 以便它断开 (图 1I, 红色虚线)。 注: 如有必要, 解剖显微镜可以帮助这一过程, 以确保仪器的准确放置和操作。</li>
	<li>使用细钳, 撬开大从颞骨的其余部分, 露出骨迷宫的耳蜗 (图 1J, 红色括号)。</li>
	<li>在解剖范围视野下放置一个足够大的培养皿以容纳颞骨 (例如, 30 毫米或更大), 并填充足够的磷酸盐缓冲盐水 (PBS) 来覆盖颞骨 (如例如, 0.5 到1毫升)。</li>
	<li>淹没世俗骨头与大去除在 PBS 与耳蜗的卵形和圆的窗口朝向。 注意: 从耳蜗中取出前庭系统是不必要的, 因为前庭系统很好地充当了操纵和定向耳蜗功能的手柄 (图 1I和图 2B)。</li>
	<li>使用超细钳, 取出镫动脉, 穿过镫。</li>
	<li>插入一个尖端的超细钳通过拱镫的地方, 动脉通过, 并微妙地解除镫向上。关节镫应该从椭圆形窗口中抬起。</li>
	<li>用一个尖端的超细钳, 刺穿椭圆形和圆形窗口。</li>
	<li>使用 1x PBS 填充28.5 口径注射器与油管 (内径0.28 毫米, 外径 0.61 mm) 连接, 定位油管在椭圆形窗口 (图 2A -b)。 注: 管子的开口可以使用超细镊子来操纵, 以确保准确的放置。<ol>
		<li>为每个耳蜗使用新鲜的注射器和导管。</li>
	</ol>
	</li>
	<li>灌注2毫升的10% 抗生素抗 (防) 和10% 青霉素-链霉素 (笔链球菌) 在 PBS 通过耳蜗超过5分钟, 以消除淋巴。灌注太快, 压力过大会损害耳蜗内的细微结构。 注: 手动驾驶注射器灌注通过耳蜗的液体是有效的, 尽管灌注泵可以作为替代品使用。<ol>
		<li>如果手动灌注, 使用一对优良的, 自闭钳, 以稳定的耳蜗 (图 2A) 在灌注过程中抓住颞骨的前庭部分。 注: 虽然不需要, 这留下双手自由处理仪器;一只手可以驱动注射器 (图 2A), 而另一个则可以适当定位油管 (图 2B)。</li>
		<li>从这一步向前, 小心避免暴露耳蜗不必要的空气。将气泡引入流体空间会堵塞流体循环, 进一步干燥会破坏耳蜗内的细微结构。</li>
	</ol>
	</li>
	<li>用细钳去除和丢弃剩余的肌肉组织和骨碎片。</li>
	<li>如有必要, 在4° c 的10% 防和10% 笔链球菌感染溶液中, 在 PBS 中贮存隔离耳蜗七天, 每48小时定期更换抗生素溶液。</li>
</ol>
2. 人工耳蜗加工
<ol>
	<li>细胞
	<ol>
		<li>对于每一个耳蜗, 填补一个20毫升玻璃闪烁瓶与1% 十二烷基硫酸钠 (SDS) 溶液在去 (DI) 水, 这是用来 decellularize 耳蜗。</li>
		<li>用剃刀刀片将塑料7毫升的转移吸管的尖端切掉, 这样开口就足够大, 足以让耳蜗通过。</li>
		<li>使用转移吸管, 轻轻地将耳蜗拉入吸管, 使其刚好通过几毫米的截边。</li>
		<li>将耳蜗排出 1% SDS 溶液填充的闪烁小瓶。</li>
		<li>在去 (DI) 水中循环2毫升的 1% SDS, 通过在闪烁瓶中使用转移吸管的耳蜗。</li>
		<li>将闪烁小瓶放入转子, 并允许闪烁瓶在室温下以 10 rpm 旋转72小时。<ol>
			<li>在72小时内每24小时更换 1% SDS 溶液。改变 SDS 时要注意不要让耳蜗暴露在空气中。</li>
		</ol>
		</li>
		<li>用 DI 水冲洗耳蜗三次, 每次30分钟。</li></ol></li></ol>在用于 SDS 费用的相同闪烁瓶中进行洗涤;在这个阶段, 耳蜗是细胞的。
	
	
	<li>脱
	<ol>
		<li>从闪烁瓶中取出剩余的 DI 水, 只留下足够的时间使耳蜗被淹没。</li>
		<li>开始脱, 增加5毫升的10% 乙酸酸 (EDTA) (0.02 米) 在 DI 水中的闪烁瓶含有的耳蜗。</li>
		<li>使用转移吸管, 通过闪烁瓶中的耳蜗在 DI 水中循环2毫升 10% EDTA。</li>
		<li>将闪烁瓶放回转子, 并允许闪烁瓶在室温下以 10 rpm 旋转72小时。在72小时内每24小时改变 10% EDTA 溶液。在改变溶液时要小心, 这样耳蜗就不会暴露在空气中。</li>
		<li>用 1x PBS 冲洗耳蜗3次, 每次2小时;在这个阶段, 耳蜗是脱的。</li>
	</ol>
	</li>
	<li>存储
	<ol>
		<li>在4° c 的 PBS 中, 在10% 防、10% 笔链球菌感染溶液中储存耳蜗 72 h。 注: 定期更换抗生素溶液可能会延长贮存期;但是, 已处理的耳蜗的扩展存储尚未在本协议中得到验证。</li>
	</ol>
	</li>

3. 采购和扩大 hWJCs
<ol>
	<li>根据以前发布的协议22隔离 hWJCs。简单地, 将脐带分成3厘米的部分。</li>
	<li>小心地用手术刀在每个脐带段上做一个浅切口, 展开并露出沃顿的脐带果冻。用镊子取出血管。</li>
	<li>用足够的 PBS (例如, 40 毫升在60毫米的培养皿中) 冲洗脐段, 以淹没组织。用手术刀切碎的组织片段。在 low-glucose Dulbecco 氏改良的鹰培养基 (DMEM) 中, 用50毫升消化培养基 (0.2% 型2胶原酶, 1% 青霉素-链霉素) 在100毫米培养皿中消化组织。</li>
	<li>在 5% CO2, 37 ° c 恒温箱中, 在一个轨道振动器上, 在 50 rpm 的夜间将组织放入消化介质中。第二天, 在 PBS 的2% 种抗生素抗中, 稀消化培养基的比例为 1:16, 在室温 (~ 27 ° c) 下离心 500 x g 的颗粒细胞, 为10分钟丢弃上清。</li>
	<li>骨髓间充质干细胞生长培养基 (MSCGM) 和重组细胞重颗粒。在组织培养中处理 T-75 烧瓶的密度为 7 x 103细胞/cm2的平板细胞。文化 HWJCs 在 MSCGM, 并扩大到通过5的实验。 注: HWJCs 可 sub-cultured 为更大的 T 型烧瓶 (如, T-150, T-300), 传代, 以提高 HWJCs 的实验产量。</li>
</ol>
4. 将 hWJCs 注入细胞耳蜗
<ol>
	<li>准备细胞耳蜗
	<ol>
		<li>使用无菌技术, 在 1x PBS 中洗涤储存的耳蜗三次, 每次30分钟。</li>
		<li>将耳蜗转化为24孔板的独立井, 并加入1毫升的37° c 5ml MSCGM。</li>
		<li>孵育耳蜗在 5% CO2, 37 ° c 细胞培养孵化器为至少 1 h 在注入细胞之前。</li>
	</ol>
	</li>
	<li>离解 hWJCs 和重
	<ol>
		<li>清洗 hWJCs 与37° c 5ml 1x PBS 两次。</li>
		<li>添加足够的37° c 5ml 胰蛋白酶与 0.5% EDTA 覆盖细胞血管表面。</li>
		<li>孵育 hWJCs 5 分钟, 在 5% CO2, 37 ° c 细胞培养孵化器。<ol>
			<li>通过倒置显微镜观察细胞, 验证90% 的细胞是否与细胞培养表面分离。 注意: 当培养容器被轻轻拍打时移动的细胞已经分离。保持静止的细胞, 没有分离。</li>
			<li>轻轻地敲击烧瓶的两侧, 以进一步去除附着的细胞。</li>
		</ol>
		</li>
		<li>采用无菌技术, 将 hWJCs 从培养容器转移到50毫升的圆锥管, 其中含有相当于 MSCGM 的体积, 以用于游离细胞的胰蛋白酶。</li>
		<li>在室温 (〜27° c) 下, 离心在 500 x g 的 hWJCs 为5分钟。</li>
		<li>吸气上清, 和重 hWJCs 在 MSCGM 的浓度为50万细胞/毫升。</li>
	</ol>
	</li>
	<li>灌注耳蜗
	<ol>
		<li>转移耳蜗到一个新的24井板使用转移吸管。</li>
		<li>在每个耳蜗上加一滴 MSCGM, 防止耳蜗干燥。</li>
		<li>用超细镊子巧妙地定位耳蜗, 使椭圆形和圆形的窗户朝上。 注意: 当耳蜗容易受损时, 直接处理耳蜗时要格外小心。</li>
		<li>使用一个无菌28.5 口径胰岛素注射器与连接的油管, 绘制0.2 毫升悬浮 hWJCs (10万细胞)。</li>
		<li>使用灭菌的细钳, 放置在椭圆形窗口的油管。</li>
		<li>微妙和缓慢灌注的耳蜗与0.2 毫升的悬浮 hWJCs 使用28.5 口径胰岛素注射器连接到聚乙烯管 (外径: 0.61 毫米, 内径: 0.28 毫米) 在大约5分钟。</li>
		<li>灌注后, 将0.8 毫升的37° c 5ml MSCGM 到含有耳蜗的井中, 使井内总容积增加到1毫升。</li>
		<li>对每一个耳蜗重复步骤 4.3. 3-4. 3.7, 每次使用新鲜的注射器和导管。</li>
		<li>放置灌注耳蜗在 5% CO2, 37 ° c 细胞培养孵化器, 并改变媒体每周三次。</li>
	</ol>
	</li>
</ol>
5. 耳蜗的采集和保存
注: 耳蜗可在任何时间点培养和收获30天 post-perfusion。
<ol>
	<li>为了保存耳蜗, 在一个摇摆平台上, 在4° c 的晚上, 在 1x PBS 中修复4% 甲醛。</li>
	<li>固定后, 用 1x PBS 清洗耳蜗三次, 每次5分钟。</li>
	<li>在石蜡包埋之前, 用乙醇和脂肪烃溶剂逐渐将耳蜗脱水。</li>
	<li>截面样品的厚度为10µm 使用切片和安装在玻璃显微镜幻灯片。 注: 样品现在可以使用标准的协议进行组织学或免疫组织化学处理。</li>
</ol>

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J., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Improving+Viability+and+Transfection+Efficiency+with+Human+Umbilical+Cord+Wharton's+Jelly+Cells+Through+Use+of+a+ROCK+Inhibitor.">Improving Viability and Transfection Efficiency with Human Umbilical Cord Wharton's Jelly Cells Through Use of a ROCK Inhibitor.</a> Cell Reprogram. , (2014).</li><li>Mellott, A. J., Shinogle, H. E., Moore, D. S., Detamore, M. S. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Fluorescent+Photo-conversion:+A+second+chance+to+label+unique+cells.">Fluorescent Photo-conversion: A second chance to label unique cells.</a> Cell Mol Bioeng. 8 (1), 187-196 (2015).</li><li>Mitchell, K. E., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Matrix+cells+from+Wharton's+jelly+form+neurons+and+glia.">Matrix cells from Wharton's jelly form neurons and glia.</a> Stem Cells. 21 (1), 50-60 (2003).</li><li>Mellott, A. J., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Nonviral+Reprogramming+of+Human+Wharton's+Jelly+Cells+Reveals+Differences+Between+ATOH1+Homologues.">Nonviral Reprogramming of Human Wharton's Jelly Cells Reveals Differences Between ATOH1 Homologues.</a> Tissue Eng Part A. 21 (11-12), 1795-1809 (2015).</li><li>Buytaert, J. A., Johnson, S. B., Dierick, M., Salih, W. H., Santi, P. A. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=MicroCT+versus+sTSLIM+3D+imaging+of+the+mouse+cochlea.">MicroCT versus sTSLIM 3D imaging of the mouse cochlea.</a> J Histochem Cytochem. 61 (5), 382-395 (2013).</li><li>Nonoyama, H., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Investigation+of+the+ototoxicity+of+gadoteridol+(ProHance)+and+gadodiamide+(Omniscan)+in+mice.">Investigation of the ototoxicity of gadoteridol (ProHance) and gadodiamide (Omniscan) in mice.</a> Acta Otolaryngol. 136 (11), 1091-1096 (2016).</li><li>Dai, C., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Rhesus+Cochlear+and+Vestibular+Functions+Are+Preserved+After+Inner+Ear+Injection+of+Saline+Volume+Sufficient+for+Gene+Therapy+Delivery.">Rhesus Cochlear and Vestibular Functions Are Preserved After Inner Ear Injection of Saline Volume Sufficient for Gene Therapy Delivery.</a> J Assoc Res Otolaryngol. , (2017).</li><li>Sutherland, A. J., Converse, G. L., Hopkins, R. A., Detamore, M. S. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=The+bioactivity+of+cartilage+extracellular+matrix+in+articular+cartilage+regeneration.">The bioactivity of cartilage extracellular matrix in articular cartilage regeneration.</a> Adv Healthc Mater. 4 (1), 29-39 (2015).</li></ol>

作者没有什么可透露的。

我们已经成功地证明, 原生耳蜗细胞可以通过细胞的过程从耳蜗中移除, 这使得耳蜗的使用成为一个复杂的三维组织支架。桑蒂et al。15开发了 decellularizing 耳蜗的初始方法, 并通过光片显微镜23精确估计了许多耳蜗结构的体积。这种早期的工作为组织工程和细胞培养技术提供了强有力的基础。细胞耳蜗可以成功地注入细胞, 然后培养一段长时间。从理论上讲...

耳蜗是一种复杂的螺旋结构, 在颞骨中发现。它是由外骨迷宫和同心, 内膜迷宫1组成。膜迷宫包括三流体空间: scala vestibuli, scala 媒体, 和 scala 索1。scala 媒介安置感觉上皮, 由许多细胞类型组成, 但感觉毛细胞 (HC), 传感器机械能在音波到神经冲动2, 是特别感兴趣的。接触到声学创伤3,4,5, 药物治疗6, 疾病7,8, 和老化9都可以通过 HC 死亡导致听觉功能受损。哺乳动物的毛细胞丢失是永久性的, 不像禽 HCs, 它可以在受伤后再生10。
各种当代研究努力试图恢复失去的 HCs, 虽然具体的实验方法有所不同。在体外分化的感觉上皮和干细胞植入的基因表达的操作是该领域的主要方法, 尽管寻求将干细胞分化成耳蜗 organoids 的诱导方法已经已尝试11,12,13。每种方法要么直接依赖干细胞, 要么是干细胞所使用的发育线索;但是, 第二个共享的和潜在的关键元素是耳蜗自身的 ECM14,15。
ECM 不仅为细胞和组织提供物理支持, 包括细胞黏附、增殖、存活和迁移的表面, 而且在 HCs 和螺旋神经节的发育中起关键作用15,16 ,17。自然发生的 ECM 提供诱导信号, 可以指导细胞表型的确定和/或细胞黏附, 增殖, 和生存18。因此, 利用细胞耳蜗结合培养 hWJCs 提供了一个独特的机会, 以探索的作用, ECM 和 HC 再生。HWJCs 是一种容易获得的, 争议细胞类型, 从人类脐带中分离出来, 表现得像间充质干细胞19。HWJCs 已经显示了区分感觉神经细胞谱系的能力20,21。因此, 目前的协议细节的隔离, 细胞, 并灌注耳蜗从 C57BL 鼠胴体与 hWJCs 的内耳组织工程。

<table>
	<tbody>
		<tr>
			<td>Allegra X-14R Centrifuge</td>
			<td>Beckman-Coulter</td>
			<td>B08861</td>
			<td></td>
		</tr>
		<tr>
			<td>Intramedic Semi-Rigid Tubing</td>
			<td>Becton Dickinson</td>
			<td>427401</td>
			<td></td>
		</tr>
		<tr>
			<td>New Brunswick Innova 2000 Orbital Shaler</td>
			<td>Eppendorf</td>
			<td>M1190-0002</td>
			<td></td>
		</tr>
		<tr>
			<td>Surgical Scissors</td>
			<td>Fine Science Tools</td>
			<td>14060-10</td>
			<td></td>
		</tr>
		<tr>
			<td>Fine Forceps</td>
			<td>Fine Science Tools</td>
			<td>11370-40</td>
			<td></td>
		</tr>
		<tr>
			<td>Ultra-Fine Forceps</td>
			<td>Fine Science Tools</td>
			<td>18155-13</td>
			<td></td>
		</tr>
		<tr>
			<td>50-mL Conical Tubes</td>
			<td>Fisher Scientific</td>
			<td>12565271</td>
			<td></td>
		</tr>
		<tr>
			<td>Petri Dish</td>
			<td>Fisher Scientific</td>
			<td>FB087579B</td>
			<td></td>
		</tr>
		<tr>
			<td>U-100 Insulin Syringe</td>
			<td>Fisher Scientific</td>
			<td>14-829-1B</td>
			<td></td>
		</tr>
		<tr>
			<td>Scintillation Vial</td>
			<td>Fisher Scientific</td>
			<td>03-341-73</td>
			<td></td>
		</tr>
		<tr>
			<td>Rotator</td>
			<td>Fisher Scientific</td>
			<td>88-861-049</td>
			<td></td>
		</tr>
		<tr>
			<td>Transfer Pipette</td>
			<td>Fisher Scientific</td>
			<td>22-170-404</td>
			<td></td>
		</tr>
		<tr>
			<td>Razor Blade</td>
			<td>Fisher Scientific</td>
			<td>12-640</td>
			<td></td>
		</tr>
		<tr>
			<td>Antibiotic-Antimycotic</td>
			<td>Fisher Scientific</td>
			<td>15-240-062</td>
			<td></td>
		</tr>
		<tr>
			<td>Penicillin-Streptomycin</td>
			<td>Fisher Scientific</td>
			<td>15-140-122</td>
			<td></td>
		</tr>
		<tr>
			<td>24-Well Plate</td>
			<td>Fisher Scientific</td>
			<td>07-200-84</td>
			<td></td>
		</tr>
		<tr>
			<td>SuperFrost PLUS Glass Microscope Slides</td>
			<td>Fisher Scientific</td>
			<td>12-550-15</td>
			<td></td>
		</tr>
		<tr>
			<td>Transfer Pipette</td>
			<td>Fisher Scientific</td>
			<td>22-170-404</td>
			<td></td>
		</tr>
		<tr>
			<td>ProLong Gold Antifade Mountant with DAPI</td>
			<td>Fisher Scientific</td>
			<td>P36935</td>
			<td></td>
		</tr>
		<tr>
			<td>Clear-Rite 3</td>
			<td>Fisher Scientific</td>
			<td>22-046341</td>
			<td></td>
		</tr>
		<tr>
			<td>Thermo Scientific Forma Series II 3110 Water-Jacekted CO2 Incubator</td>
			<td>Fisher Scientific</td>
			<td>13-998-078</td>
			<td></td>
		</tr>
		<tr>
			<td>Mesenchymal Stem Cell Growth Medium</td>
			<td>Lonza</td>
			<td>PT-3001</td>
			<td></td>
		</tr>
		<tr>
			<td>Trypsin-EDTA</td>
			<td>Lonza</td>
			<td>CC-3232</td>
			<td></td>
		</tr>
		<tr>
			<td>TPP T-75 Culture Flask</td>
			<td>MidSci</td>
			<td>TP90076</td>
			<td></td>
		</tr>
		<tr>
			<td>TPP T-150 Culture Flask</td>
			<td>MidSci</td>
			<td>TP90151</td>
			<td></td>
		</tr>
		<tr>
			<td>TPP T-300 Culture Flask</td>
			<td>MidSci</td>
			<td>TP90301</td>
			<td></td>
		</tr>
		<tr>
			<td>Dissection Microscope</td>
			<td>Nikon Instruments</td>
			<td>SMZ800</td>
			<td></td>
		</tr>
		<tr>
			<td>Nikon Eclipse Ts2R-FL Inverted Microscope</td>
			<td>Nikon Instruments</td>
			<td>MFA51010</td>
			<td></td>
		</tr>
		<tr>
			<td>NuAire Class II, Type A2 Biosafety Cabinet</td>
			<td>NuAire</td>
			<td>NU-425-600</td>
			<td></td>
		</tr>
		<tr>
			<td>1X PBS</td>
			<td>Sigma-Aldrich</td>
			<td>P5368-10PAK</td>
			<td></td>
		</tr>
		<tr>
			<td>1% SDS Solution</td>
			<td>Sigma-Aldrich</td>
			<td>436143-100G</td>
			<td></td>
		</tr>
		<tr>
			<td>10% EDTA</td>
			<td>Sigma-Aldrich</td>
			<td>E9884-100G</td>
			<td></td>
		</tr>
	</tbody>
</table>

a protocol for decellularizing mouse cochleae for inner ear tissue engineering

在哺乳动物, mechanosensory 的毛细胞, 促进听力缺乏再生能力, 这是有限的治疗听力损失。目前的再生医学战略的重点是移植干细胞或在内耳周围的支持细胞的遗传操作, 以鼓励更换受损的干细胞, 以纠正听力损失。然而, 细胞外基质 (ECM) 在诱导和维持毛细胞功能方面可能起着至关重要的作用, 而且尚未得到很好的研究。利用耳蜗 ECM 作为培养成体干细胞的支架, 可以为细胞外环境的组成和结构对维持听觉功能的支持提供独特的见解。在这里, 我们提出了一种方法, 分离和 decellularizing 耳蜗从小鼠使用作为支架接受灌注成体干细胞。在当前的协议中, 耳蜗是从被安乐死的小鼠、细胞和脱中分离出来的。之后, 人沃顿的果冻细胞 (hWJCs), 从脐带被隔离, 仔细地灌注到每个耳蜗。耳蜗被用作生物反应器, 细胞被培养30天后进行分析处理。细胞耳蜗保留可辨认的胞外结构, 但没有揭示细胞的存在或 DNA 的明显片段。耳蜗内灌注的细胞侵入耳蜗的大部分内部和外部, 在30天的时间内没有发生任何事件。因此, 目前的方法可以用来研究如何耳蜗 ECM 影响细胞的发展和行为。

使用这里提出的方法, 成功的细胞耳蜗评估通过检查是否存在或没有 DNA 通过 4, 6-diamidino-2-吲 (DAPI) 染色。如果在细胞耳蜗内没有发现 DNA, 耳蜗被认为是完全细胞的。以前的实验中, 没有经过细胞或脱的原生耳蜗被用作一种阳性对照, 用以说明传统上在 C57BL 鼠耳蜗中发现的结构和细胞。细胞-脱耳蜗, 没有任何 hWJCs 注入, 被用作负控制。在组织切片的任何区域 (...

Watch this Scientific Journal Video about A Protocol for Decellularizing Mouse Cochleae for Inner Ear Tissue Engineering at JoVE.com

A Protocol for Decellularizing Mouse Cochleae for Inner Ear Tissue Engineering

用于内耳组织工程的 Decellularizing 小鼠耳蜗协议

本协议的目的是演示一种有效的方法, decellularize 和 decalcify 小鼠耳蜗作为组织工程应用的支架使用。

In mammals, mechanosensory hair cells that facilitate hearing lack the ability to regenerate, which has limited treatments for hearing loss. Current ...

a-protocol-for-decellularizing-mouse-cochleae-for-inner-ear-tissue

Research

JoVE Journal

Bioengineering

10.7K 次观看.  University of Kansas Medical Center. 本协议的目的是演示一种有效的方法, decellularize 和 decalcify 小鼠耳蜗作为组织工程应用的支架使用。

视频: A Protocol for Decellularizing Mouse Cochleae for Inner Ear Tissue Engineering

Fabrication of Biologically Derived Injectable Materials for Myocardial Tissue Engineering

This protocol provides methods for the preparation of an injectable extracellular matrix (ECM) gel for myocardial tissue engineering applications. Briefly, decellularized tissue is lyophilized, milled, enzymatically digested, and then brought to physiological pH. The lyophilization removes all water content from the tissue, resulting in dry ECM that can be ground into a fine powder with a small mill. After milling, the ECM powder is digested with pepsin to form an injectable matrix. After adjustment to pH 7.4, the liquid matrix material can be injected into the myocardium. Results of previous characterization assays have shown that matrix gels produced from decellularized pericardial and myocardial tissue retain native ECM components, including diverse proteins, peptides and glycosaminoglycans. Given the use of this material for tissue engineering, in vivo characterization is especially useful; here, a method for performing an intramural injection into the left ventricular (LV) free wall is presented as a means of analyzing the host response to the matrix gel in a small animal model. Access to the chest cavity is gained through the diaphragm and the injection is made slightly above the apex in the LV free wall. The biologically derived scaffold can be visualized by biotin-labeling before injection and then staining tissue sections with a horse radish peroxidase-conjugated neutravidin and visualizing via diaminobenzidine (DAB) staining. Analysis of the injection region can also be done with histological and immunohistochemical staining. In this way, the previously examined pericardial and myocardial matrix gels were shown to form fibrous, porous networks and promote vessel formation within the injection region.

Methods for preparing an injectable matrix gel from decellularized tissue and injecting it into rat myocardium in vivo are described.

心肌组织工程生物衍生的注射材料的制备

This protocol provides methods for the preparation of an injectable extracellular matrix (ECM) gel for myocardial tissue engineering applications.  ...

科研

生物工程

Procedure for Lung Engineering

Lung tissue, including lung cancer and chronic lung diseases such as chronic obstructive pulmonary disease, cumulatively account for some 280,000 deaths annually; chronic obstructive pulmonary disease is currently the fourth leading cause of death in the United States1. Contributing to this mortality is the fact that lungs do not generally repair or regenerate beyond the microscopic, cellular level. Therefore, lung tissue that is damaged by degeneration or infection, or lung tissue that is surgically resected is not functionally replaced in vivo. To explore whether lung tissue can be generated in vitro, we treated lungs from adult rats using a procedure that removes cellular components to produce an acellular lung extracellular matrix scaffold. This scaffold retains the hierarchical branching structures of airways and vasculature, as well as a largely intact basement membrane, which comprises collagen IV, laminin, and fibronectin. The scaffold is mounted in a bioreactor designed to mimic critical aspects of lung physiology, such as negative pressure ventilation and pulsatile vascular perfusion. By culturing pulmonary epithelium and vascular endothelium within the bioreactor-mounted scaffold, we are able to generate lung tissue that is phenotypically comparable to native lung tissue and that is able to participate in gas exchange for short time intervals (45-120 minutes). These results are encouraging, and suggest that repopulation of lung matrix is a viable strategy for lung regeneration. This possibility presents an opportunity not only to work toward increasing the supply of lung tissue for transplantation, but also to study respiratory cell and molecular biology in vitro for longer time periods and in a more accurate microenvironment than has previously been possible.

We have developed a decellularized lung extracellular matrix and novel biomimetic bioreactor that can be used to generate functional lung tissue. By seeding cells into the matrix and culturing in the bioreactor, we generate tissue that demonstrates effective gas exchange when transplanted in vivo for short periods of time.

龙工程的程序

Lung tissue, including lung cancer and chronic lung diseases such as chronic obstructive pulmonary disease, cumulatively account for some 280,000 deaths ...

Elastomeric PGS Scaffolds in Arterial Tissue Engineering

Cardiovascular disease is one of the leading cause of mortality in the US and especially, coronary artery disease increases with an aging population and increasing obesity1. Currently, bypass surgery using autologous vessels, allografts, and synthetic grafts are known as a commonly used for arterial substitutes2. However, these grafts have limited applications when an inner diameter of arteries is less than 6 mm due to low availability, thrombotic complications, compliance mismatch, and late intimal hyperplasia3,4. To overcome these limitations, tissue engineering has been successfully applied as a promising alternative to develop small-diameter arterial constructs that are nonthrombogenic, robust, and compliant. Several previous studies have developed small-diameter arterial constructs with tri-lamellar structure, excellent mechanical properties and burst pressure comparable to native arteries5,6. While high tensile strength and burst pressure by increasing collagen production from a rigid material or cell sheet scaffold, these constructs still had low elastin production and compliance, which is a major problem to cause graft failure after implantation. Considering these issues, we hypothesized that an elastometric biomaterial combined with mechanical conditioning would provide elasticity and conduct mechanical signals more efficiently to vascular cells, which increase extracellular matrix production and support cellular orientation.
The objective of this report is to introduce a fabrication technique of porous tubular scaffolds and a dynamic mechanical conditioning for applying them to arterial tissue engineering. We used a biodegradable elastomer, poly (glycerol sebacate) (PGS)7 for fabricating porous tubular scaffolds from the salt fusion method. Adult primary baboon smooth muscle cells (SMCs) were seeded on the lumen of scaffolds, which cultured in our designed pulsatile flow bioreactor for 3 weeks. PGS scaffolds had consistent thickness and randomly distributed macro- and micro-pores. Mechanical conditioning from pulsatile flow bioreactor supported SMC orientation and enhanced ECM production in scaffolds. These results suggest that elastomeric scaffolds and mechanical conditioning of bioreactor culture may be a promising method for arterial tissue engineering.

Elastomeric PGS scaffolds with vascular smooth muscle cells cultured in a pulsatile flow bioreactor may lead to promising small-diameter arterial constructs with native ECM production in a relatively short culture period.

弹性PGS的动脉组织工程支架

Cardiovascular disease is one of the leading cause of mortality in the US and especially, coronary artery disease increases with an aging population and ...

Encapsulation of Cardiomyocytes in a Fibrin Hydrogel for Cardiac Tissue Engineering

Culturing cells in a three dimensional hydrogel environment is an important technique for developing constructs for tissue engineering as well as studying cellular responses under various culture conditions in vitro. The three dimensional environment more closely mimics what the cells observe in vivo due to the application of mechanical and chemical stimuli in all dimensions 1. Three-dimensional hydrogels can either be made from synthetic polymers such as PEG-DA 2 and PLGA 3 or a number of naturally occurring proteins such as collagen 4, hyaluronic acid 5 or fibrin 6,7. Hydrogels created from fibrin, a naturally occurring blood clotting protein, can polymerize to form a mesh that is part of the body's natural wound healing processes 8. Fibrin is cell-degradable and potentially autologous 9, making it an ideal temporary scaffold for tissue engineering.
Here we describe in detail the isolation of neonatal cardiomyocytes from three day old rat pups and the preparation of the cells for encapsulation in fibrin hydrogel constructs for tissue engineering. Neonatal myocytes are a common cell source used for in vitro studies in cardiac tissue formation and engineering 4. Fibrin gel is created by mixing fibrinogen with the enzyme thrombin. Thrombin cleaves fibrinopeptides FpA and FpB from fibrinogen, revealing binding sites that interact with other monomers 10. These interactions cause the monomers to self-assemble into fibers that form the hydrogel mesh. Because the timing of this enzymatic reaction can be adjusted by altering the ratio of thrombin to fibrinogen, or the ratio of calcium to thrombin, one can injection mold constructs with a number of different geometries 11,12. Further we can generate alignment of the resulting tissue by how we constrain the gel during culture 13.
After culturing the engineered cardiac tissue constructs for two weeks under static conditions, the cardiac cells have begun to remodel the construct and can generate a contraction force under electrical pacing conditions 6. As part of this protocol, we also describe methods for analyzing the tissue engineered myocardium after the culture period including functional analysis of the active force generated by the cardiac muscle construct upon electrical stimulation, as well as methods for determining final cell viability (Live-Dead assay) and immunohistological staining to examine the expression and morphology of typical proteins important for contraction (Myosin Heavy Chain or MHC) and cellular coupling (Connexin 43 or Cx43) between myocytes.

We describe the isolation of neonatal cardiomyocytes and the preparation of the cells for encapsulation in fibrin hydrogel constructs for tissue engineering. We describe methods for analyzing the tissue engineered myocardium after the culture period including active force generated upon electrical stimulation and cell viability and immunohistological staining.

心肌在心脏组织工程中的纤维蛋白凝胶的封装

Culturing cells in a three dimensional hydrogel environment is an important technique for developing constructs for tissue engineering as well as studying ...

Directed Cellular Self-Assembly to Fabricate Cell-Derived Tissue Rings for Biomechanical Analysis and Tissue Engineering

Each year, hundreds of thousands of patients undergo coronary artery bypass surgery in the United States.1 Approximately one third of these patients do not have suitable autologous donor vessels due to disease progression or previous harvest. The aim of vascular tissue engineering is to develop a suitable alternative source for these bypass grafts. In addition, engineered vascular tissue may prove valuable as living vascular models to study cardiovascular diseases. Several promising approaches to engineering blood vessels have been explored, with many recent studies focusing on development and analysis of cell-based methods.2-5 Herein, we present a method to rapidly self-assemble cells into 3D tissue rings that can be used in vitro to model vascular tissues.
To do this, suspensions of smooth muscle cells are seeded into round-bottomed annular agarose wells. The non-adhesive properties of the agarose allow the cells to settle, aggregate and contract around a post at the center of the well to form a cohesive tissue ring.6,7 These rings can be cultured for several days prior to harvesting for mechanical, physiological, biochemical, or histological analysis. We have shown that these cell-derived tissue rings yield at 100-500 kPa ultimate tensile strength8 which exceeds the value reported for other tissue engineered vascular constructs cultured for similar durations (&lt;30 kPa).9,10 Our results demonstrate that robust cell-derived vascular tissue ring generation can be achieved within a short time period, and offers the opportunity for direct and quantitative assessment of the contributions of cells and cell-derived matrix (CDM) to vascular tissue structure and function.

This article outlines a versatile method to create cell-derived tissue rings by cellular self-assembly. Smooth muscle cells seeded into ring-shaped agarose wells aggregate and contract to form robust three-dimensional (3D) tissues within 7 days. Millimeter-scale tissue rings are conducive to mechanical testing and serve as building blocks for tissue assembly.

定向细胞自我组装制造，为生物力学分析与组织工程细胞衍生的组织环

Each year, hundreds of thousands of patients undergo coronary artery bypass surgery in the United States.1 Approximately one third of these patients do ...

Postproduction Processing of Electrospun Fibres for Tissue Engineering

Electrospinning is a commonly used and versatile method to produce scaffolds (often biodegradable) for 3D tissue engineering.1, 2, 3 Many tissues in vivo undergo biaxial distension to varying extents such as skin, bladder, pelvic floor and even the hard palate as children grow. In producing scaffolds for these purposes there is a need to develop scaffolds of appropriate biomechanical properties (whether achieved without or with cells) and which are sterile for clinical use. The focus of this paper is not how to establish basic electrospinning parameters (as there is extensive literature on electrospinning) but on how to modify spun scaffolds post production to make them fit for tissue engineering purposes - here thickness, mechanical properties and sterilisation (required for clinical use) are considered and we also describe how cells can be cultured on scaffolds and subjected to biaxial strain to condition them for specific applications.
Electrospinning tends to produce thin sheets; as the electrospinning collector becomes coated with insulating fibres it becomes a poor conductor such that fibres no longer deposit on it. Hence we describe approaches to produce thicker structures by heat or vapour annealing increasing the strength of scaffolds but not necessarily the elasticity. Sequential spinning of scaffolds of different polymers to achieve complex scaffolds is also described. Sterilisation methodologies can adversely affect strength and elasticity of scaffolds. We compare three methods for their effects on the biomechanical properties on electrospun scaffolds of poly lactic-co-glycolic acid (PLGA).
Imaging of cells on scaffolds and assessment of production of extracellular matrix (ECM) proteins by cells on scaffolds is described. Culturing cells on scaffolds in vitro can improve scaffold strength and elasticity but the tissue engineering literature shows that cells often fail to produce appropriate ECM when cultured under static conditions. There are few commercial systems available that allow one to culture cells on scaffolds under dynamic conditioning regimes - one example is the Bose Electroforce 3100 which can be used to exert a conditioning programme on cells in scaffolds held using mechanical grips within a media filled chamber.4 An approach to a budget cell culture bioreactor for controlled distortion in 2 dimensions is described. We show that cells can be induced to produce elastin under these conditions. Finally assessment of the biomechanical properties of processed scaffolds cultured with or without cells is described.

Electrospun scaffolds can be processed post production for tissue engineering applications. Here we describe methods for spinning complex scaffolds (by consecutive spinning), for making thicker scaffolds (by multi-layering using heat or vapour annealing), for achieving sterility (aseptic production or sterilisation post production) and for achieving appropriate biomechanical properties.

静电纤维组织工程的后期处理

Electrospinning is a commonly used and versatile method to produce  scaffolds (often biodegradable) for 3D tissue engineering.1, 2, 3 Many tissues in vivo ...

Capillary Force Lithography for Cardiac Tissue Engineering

Cardiovascular disease remains the leading cause of death worldwide1. Cardiac tissue engineering holds much promise to deliver groundbreaking medical discoveries with the aims of developing functional tissues for cardiac regeneration as well as in vitro screening assays. However, the ability to create high-fidelity models of heart tissue has proven difficult. The heart&#8217;s extracellular matrix (ECM) is a complex structure consisting of both biochemical and biomechanical signals ranging from the micro- to the nanometer scale2. Local mechanical loading conditions and cell-ECM interactions have recently been recognized as vital components in cardiac tissue engineering3-5.
A large portion of the cardiac ECM is composed of aligned collagen fibers with nano-scale diameters that significantly influences tissue architecture and electromechanical coupling2. Unfortunately, few methods have been able to mimic the organization of ECM fibers down to the nanometer scale. Recent advancements in nanofabrication techniques, however, have enabled the design and fabrication of scalable scaffolds that mimic the in vivo structural and substrate stiffness cues of the ECM in the heart6-9.
Here we present the development of two reproducible, cost-effective, and scalable nanopatterning processes for the functional alignment of cardiac cells using the biocompatible polymer poly(lactide-co-glycolide) (PLGA)8 and a polyurethane (PU) based polymer. These anisotropically nanofabricated substrata (ANFS) mimic the underlying ECM of well-organized, aligned tissues and can be used to investigate the role of nanotopography on cell morphology and function10-14.
Using a nanopatterned (NP) silicon master as a template, a polyurethane acrylate (PUA) mold is fabricated. This PUA mold is then used to pattern the PU or PLGA hydrogel via UV-assisted or solvent-mediated capillary force lithography (CFL), respectively15,16. Briefly, PU or PLGA pre-polymer is drop dispensed onto a glass coverslip and the PUA mold is placed on top. For UV-assisted CFL, the PU is then exposed to UV radiation (&#955; = 250-400 nm) for curing. For solvent-mediated CFL, the PLGA is embossed using heat (120 &#176;C) and pressure (100 kPa). After curing, the PUA mold is peeled off, leaving behind an ANFS for cell culture. Primary cells, such as neonatal rat ventricular myocytes, as well as human pluripotent stem cell-derived cardiomyocytes, can be maintained on the ANFS2.

In this protocol, we demonstrate the fabrication of biomimetic cardiac cell culture substrata made from two distinct polymeric materials using capillary force lithography. The described methods provide a scalable, cost-effective technique to engineer the structure and function of macroscopic cardiac tissues for in vitro and in vivo applications.

毛细力光刻技术在心脏组织工程

Cardiovascular disease remains the leading cause of death worldwide1. Cardiac tissue engineering holds much promise to deliver groundbreaking medical ...

Protocol for Relative Hydrodynamic Assessment of Tri-leaflet Polymer Valves

Limitations of currently available prosthetic valves, xenografts, and homografts have prompted a recent resurgence of developments in the area of tri-leaflet polymer valve prostheses. However, identification of a protocol for initial assessment of polymer valve hydrodynamic functionality is paramount during the early stages of the design process. Traditional in vitro pulse duplicator systems are not configured to accommodate flexible tri-leaflet materials; in addition, assessment of polymer valve functionality needs to be made in a relative context to native and prosthetic heart valves under identical test conditions so that variability in measurements from different instruments can be avoided. Accordingly, we conducted hydrodynamic assessment of i) native (n = 4, mean diameter, D = 20 mm), ii) bi-leaflet mechanical (n= 2, D = 23 mm) and iii) polymer valves (n = 5, D = 22 mm) via the use of a commercially available pulse duplicator system (ViVitro Labs Inc, Victoria, BC) that was modified to accommodate tri-leaflet valve geometries. Tri-leaflet silicone valves developed at the University of Florida comprised the polymer valve group. A mixture in the ratio of 35:65 glycerin to water was used to mimic blood physical properties. Instantaneous flow rate was measured at the interface of the left ventricle and aortic units while pressure was recorded at the ventricular and aortic positions. Bi-leaflet and native valve data from the literature was used to validate flow and pressure readings. The following hydrodynamic metrics were reported: forward flow pressure drop, aortic root mean square forward flow rate, aortic closing, leakage&#160;and regurgitant volume, transaortic closing, leakage, and total energy losses. Representative results indicated that hydrodynamic metrics from the three valve groups could be successfully obtained by incorporating a custom-built assembly into a commercially available pulse duplicator system and subsequently, objectively compared to provide insights on functional aspects of polymer valve design.

There has been renewed interest in developing polymer valves. Here, the objectives are to demonstrate the feasibility of modifying a commercial pulse duplicator to accommodate tri-leaflet geometries and to define a protocol to present polymer valve hydrodynamic data in comparison to native and prosthetic valve data collected under near-identical conditions.

协议相对液力评估的三单张聚合物阀

Limitations of currently available prosthetic valves, xenografts, and homografts have prompted a recent resurgence of developments in the area of ...

Design of a Biaxial Mechanical Loading Bioreactor for Tissue Engineering

We designed a loading device that is capable of applying uniaxial or biaxial mechanical strain to a tissue engineered biocomposites fabricated for transplantation. While the device primarily functions as a bioreactor that mimics the native mechanical strains, it is also outfitted with a load cell for providing force feedback or mechanical testing of the constructs. The device subjects engineered cartilage constructs to biaxial mechanical loading with great precision of loading dose (amplitude and frequency) and is compact enough to fit inside a standard tissue culture incubator. It loads samples directly in a tissue culture plate, and multiple plate sizes are compatible with the system. The device has been designed using components manufactured for precision-guided laser applications. Bi-axial loading is accomplished by two orthogonal stages. The stages have a 50 mm travel range and are driven independently by stepper motor actuators, controlled by a closed-loop stepper motor driver that features micro-stepping capabilities, enabling step sizes of less than 50 nm. A polysulfone loading platen is coupled to the bi-axial moving platform. Movements of the stages are controlled by Thor-labs Advanced Positioning Technology (APT) software. The stepper motor driver is used with the software to adjust load parameters of frequency and amplitude of both shear and compression independently and simultaneously. Positional feedback is provided by linear optical encoders that have a bidirectional repeatability of 0.1 &mu;m and a resolution of 20 nm, translating to a positional accuracy of less than 3 &mu;m over the full 50 mm of travel. These encoders provide the necessary position feedback to the drive electronics to ensure true nanopositioning capabilities. In order to provide the force feedback to detect contact and evaluate loading responses, a precision miniature load cell is positioned between the loading platen and the moving platform. The load cell has high accuracies of 0.15% to 0.25% full scale.

We designed a novel mechanical loading bioreactor that can apply uniaxial or biaxial mechanical strain to a cartilage biocomposite prior to transplantation into an articular cartilage defect.

组织工程生物反应器的双轴机械装载设计

We  designed a loading device that is capable of applying uniaxial or biaxial  mechanical strain to a tissue engineered biocomposites fabricated for  ...

Engineering 3D Cellularized Collagen Gels for Vascular Tissue Regeneration

Synthetic materials are known to initiate clinical complications such as inflammation, stenosis, and infections when implanted as vascular substitutes. Collagen has been extensively used for a wide range of biomedical applications and is considered a valid alternative to synthetic materials due to its inherent biocompatibility (i.e., low antigenicity, inflammation, and cytotoxic responses). However, the limited mechanical properties and the related low hand-ability of collagen gels have hampered their use as scaffold materials for vascular tissue engineering. Therefore, the rationale behind this work was first to engineer cellularized collagen gels into a tubular-shaped geometry and second to enhance smooth muscle cells driven reorganization of collagen matrix to obtain tissues stiff enough to be handled.
The strategy described here is based on the direct assembling of collagen and smooth muscle cells (construct) in a 3D cylindrical geometry with the use of a molding technique. This process requires a maturation period, during which the constructs are cultured in a bioreactor under static conditions (without applied external dynamic mechanical constraints) for 1 or 2 weeks. The &#8220;static bioreactor&#8221; provides a monitored and controlled sterile environment (pH, temperature, gas exchange, nutrient supply and waste removal) to the constructs. During culture period, thickness measurements were performed to evaluate the cells-driven remodeling of the collagen matrix, and glucose consumption and lactate production rates were measured to monitor the cells metabolic activity. Finally, mechanical and viscoelastic properties were assessed for the resulting tubular constructs. To this end, specific protocols and a focused know-how (manipulation, gripping, working in hydrated environment, and so on) were developed to characterize the engineered tissues.

In this work, we present a technique for the rapid fabrication of living vascular tissues by direct culturing of collagen, smooth muscle cells and endothelial cells. In addition, a new protocol for the mechanical characterization of engineered vascular tissues is described.

工程三维细胞化胶原凝胶的血管组织再生

Synthetic materials are known to initiate clinical complications such as inflammation, stenosis, and infections when implanted as vascular substitutes. ...