現在のプロジェクトは、概念資金のカンザス大学証明書によって賄われていた。我々 感謝したい看護スタッフ院長 (カンザスシティ、カンザス) で取得の蝸牛文化を支援するためひと臍帯やデビッド ・ ヨルゲンセンのことで私たちを支援のため。

カンザス大学医療センター (院長) で承認された機関動物ケアおよび使用委員会 (IACUC) のプロトコル (御急ぎでなかったら、#2014-2234) によると、動物の安楽死を含めたすべての処理を行った。
注: HWJCs は、インフォームド コンセントを提供される患者が寄贈したひと臍帯から分離した、標本 (区 IRB #15402) カンザス大学人間科目委員会によって承認されたプロトコルに従って使用されました。
1. 側頭骨収穫と蝸牛の分離
<ol>
	<li>外科はさみの鋭いペアで頭蓋骨と第一頚椎の間に切断によって適切に安楽死させた、15 週齢マウスの首をはねるし、(図 1 a) を消毒する 70% エタノールで頭をスプレーします。</li>
	<li>外科はさみ (図 1 bの-C破線) の同じシャープな対を用いた正中面で頭蓋骨を二等分します。</li>
	<li>鉗子 (図 1の矢印) のペアを使用して頭蓋骨から脳組織を削除します。</li>
	<li>存在の聴覚と前庭神経根 (図 1E-F矢印) 頭蓋骨と (図の矢印 1f) 外から外耳道の奥の頭骨を識別します。頭蓋骨 (図 1) の残りの部分から側頭骨を分離する頭蓋骨を慎重にカットします。 注: いくつかのインスタンス可能性があります可能である繊細な加工を必要とせず頭骨から頭蓋骨の周囲の骨を詮索します。</li>
	<li>ヒントは蝸牛に穴をあけるリスクはないので、外耳道 (図 1 H)、約 5 mm の開口部に微細鉗子を挿入します。優しくこじ開ける胞 (図 1 H、トップの赤斜線領域) の骨骨折 (1I を図、赤い点線線) ので。 注: 必要に応じて、解剖顕微鏡役立つ可能性がありますこのプロセスの正確な配置や楽器の操作を確保するため。</li>
	<li>微細鉗子を使用すると、(図 1 j, 赤いブラケット) 蝸牛の骨迷路の公開側頭骨からブラの残りの部分で詮索離れて。</li>
	<li>シャーレの頭骨を保持するために十分な大きさの場所 (例えば、30 mm 以上)、解離性の下にあるビューのフィールドを対象範囲し、満タンで十分なリン酸緩衝生理食塩水 (PBS) 側頭骨 (例えば1、0.5 mL) をカバーします。</li>
	<li>ブラ直面している蝸牛の楕円形や円形の窓が PBS で削除と側頭骨が水没します。 注: 前庭系を操作して定位 (1I を図と図 2 b) 蝸牛の機能のためのハンドルとしても優れていますと、蝸牛から前庭系の除去する、必要はありません。</li>
	<li>鉗子の超微細のペアを使用して、あぶみ骨を通過するアブミ骨動脈を削除します。</li>
	<li>動脈渡される、アブミ骨のアーチを超微細鉗子の先端を挿入し、繊細な上向きあぶみ骨を持ち上げます。縫合のアブミ骨は卵円窓から持ち上げる必要があります。</li>
	<li>超微細鉗子の先端を使用して、楕円の穴をあけるし、windows のラウンドします。</li>
	<li>28.5 ゲージ注射器とチューブを充填 1 × PBS を使用する (内側直径 0.28 mm、外径 0.61 mm) 管の楕円形の窓を置きます、接続されている (図 2 a -B)。 注: 管の開口部は、正確な配置を確保するため支援超微細鉗子を使用して操作できます。<ol>
		<li>各蝸牛に新鮮な注射器とチューブを使用します。</li>
	</ol>
	</li>
	<li>2 mL の 10% 抗生物質抗真菌薬 (対策) と 10% ペニシリン-ストレプトマイシン (ペン-連鎖球菌) PBS で蝸牛の 5 分以上を外リンパを削除するを灌流します。あまりにも多くの圧力で速すぎる灌蝸牛内微細構造に損傷を与えます。 注: 蝸牛を介して流体を灌流して手で注射器を手動で運転が効果的な灌流ポンプを代替として使用できます。<ol>
		<li>灌手動で場合は、側頭骨の前庭部分をつかんで灌蝸牛 (図 2 a) を安定させるために鉗子を自己終了の罰金ペアを使用します。 注意: 必須ではありませんが、これは両手を自由に楽器を処理を残します1 つの手はチューブを配置できる他、注射器 (図 2 a) を運転できる適切に (図 2 b)。</li>
		<li>このステップから転送、空気に不必要に蝸牛を公開しないように注意してください。流体の循環をブロック流体の空間に空気の泡を導入、蝸牛内の微細損傷さらに乾燥します。</li>
	</ol>
	</li>
	<li>削除し、残った微細鉗子を用いた筋肉組織や骨フラグメントを破棄します。</li>
	<li>必要に応じて、抗生物質溶液ごと 48 時間の定期的な変更で 7 日まで 10% 対策で PBS で 10% ペン連鎖球菌性ソリューション 4 ° C で分離の蝸牛を格納します。</li>
</ol>
2. 蝸牛処理
<ol>
	<li>Decellularization
	<ol>
		<li>各蝸牛の塗りつぶし 1% ナトリウム ドデシルと 20 mL ガラス シンチレーション バイアルの硫酸塩 (SDS) 水溶液のイオン (DI) の水があり、蝸牛を decellularize するために使用。</li>
		<li>開口部は十分な大きさを通過する蝸牛のかみそりの刃を使用してプラスチック製の 7 mL 転送ピペットの先端をカットします。</li>
		<li>転送ピペットを使用して、ゆっくりと引いて、蝸牛ピペットにのでそれは数ミリでカットオフ端を渡します。</li>
		<li>1 %sds ソリューションいっぱいシンチレーション バイアルに蝸牛を追放します。</li>
		<li>転送ピペットを用いたシンチレーション バイアルに蝸牛 (DI) 脱イオン水に 1 %sds の 2 mL を循環します。</li>
		<li>、ローテーターにシンチレーション バイアルを置き、部屋の温度で 72 時間 10 rpm で回転するシンチレーション バイアルを許可します。<ol>
			<li>72 時間の期間にわたって、1 %sds ソリューションすべての 24 h を変更します。決して空気に蝸牛を公開する SDS を変更するとき注意してください。</li>
		</ol>
		</li>
		<li>3 回の DI 水で各 30 分の蝸牛を洗います。</li></ol></li></ol>SDS 料金; 同じシンチレーション バイアルの洗浄を行うこの段階では、蝸牛が脱です。
	
	
	<li>脱灰
	<ol>
		<li>蝸牛を維持するのに十分なだけを残して水没、シンチレーション バイアルから残りの DI 水を削除します。</li>
		<li>脱灰を開始するには、蝸牛を含むシンチレーション バイアルに DI 水で 10% エチレンジアミン四酢酸 (EDTA) (0.02 M) 5 mL を追加します。</li>
		<li>転送ピペットを用いたシンチレーション バイアルに蝸牛を純水で 10 %edta の 2 mL を循環します。</li>
		<li>回転子に戻るシンチレーション バイアルを置き、部屋の温度で 72 時間 10 rpm で回転するシンチレーション バイアルを許可します。72 時間の期間にわたって 10 %edta 溶液のすべての 24 h を変更します。蝸牛が空気にさらされないようにソリューションを変更するときに注意します。</li>
		<li>すすぎ 3 回 1 × PBS です。 で、2 h の蝸牛この段階では、蝸牛が脱灰します。</li>
	</ol>
	</li>
	<li>ストレージ
	<ol>
		<li>72 h までの蝸牛溶液中で保存、10%、反反 10% ペン連鎖球菌 PBS で 4 ° C で 注: ストレージ期間を長く可能性があります抗生物質溶液; の定期的な変更が可能ただし、加工の蝸牛の拡張記憶域は、このプロトコルでは検証されていません。</li>
	</ol>
	</li>

3. 調達と hWJCs の拡大
<ol>
	<li>以前に発行されたプロトコル22によると hWJCs を分離します。簡単に言えば、3 cm のセクションにへその緒を分割します。</li>
	<li>慎重には、メスで浅い切開縦各臍帯のセグメントを広げるし、臍帯ワルトンを公開します。血管を削除するのに鉗子を使用します。</li>
	<li>組織が水没する (例えば、40 mL で 60 mm のペトリ皿) 十分な PBS で 2 回臍帯セグメントを洗います。メスで組織のセグメントを細かくみじん切り。100 mm ペトリ皿消化中の 50 mL の組織を消化 (0.2 %2 型コラゲナーゼ、低グルコース環境における 1% ペニシリン-ストレプトマイシン、ダルベッコ変更されたワシの媒体 (DMEM))。</li>
	<li>5% CO2、37 ° C のインキュベーターで一晩 50 回転軌道シェーカーの中の消化に組織を配置します。次の日、1:16 pbs は、抗生物質抗真菌 2% の割合で消化中を希釈し、室温で 500 × g で遠心分離によって細胞をペレット (~ 27 ° C) 10 分破棄上清の。</li>
	<li>細胞と間葉系幹細胞成長培地 (MSCGM) にペレットを再懸濁します。プレート細胞培養の 7 x 103セル/cm2の密度では、T-75 フラスコを扱われます。文化の MSCGM、HWJCs、通路 5 実験用に展開します。 注: HWJCs がサブ培養される大きい T フラスコに (例えば、T-150、T-300) 継 hWJCs 実験のための収穫を増加するとき。</li>
</ol>
4. 脱蝸牛への hWJCs の注入
<ol>
	<li>脱の蝸牛を準備します。
	<ol>
		<li>無菌技術を使用して、3 回の 30 分の 1 × PBS でストアドの蝸牛を洗浄します。</li>
		<li>24 ウェル プレートの別の井戸に蝸牛を転送し、37 ° C の 1 つの mL を追加 MSCGM で加温します。</li>
		<li>蝸牛 5% CO2セルの注入の前に 1 時間の最小値の 37 ° C 細胞文化インキュベーターで孵化させなさい。</li>
	</ol>
	</li>
	<li>HWJCs を切り離して、再懸濁します
	<ol>
		<li>37 ° C と洗浄 hWJCs 前回 1x PBS を温めた。</li>
		<li>十分な 37 ° C を追加セル容器表面に 0.5 %edta でトリプシンで加温します。</li>
		<li>HWJCs 5% CO2、37 ° C 細胞文化のインキュベーターで最大 5 分間インキュベートします。<ol>
			<li>細胞の 90% が倒立顕微鏡の下のセルを表示することによって表層文化から切り離されることを確認します。 注: 移動文化容器を軽くタップすると、細胞をデタッチします。静止したとき、細胞が分離されていません。</li>
			<li>さらに添付された細胞を取り除くためフラスコの側面が軽きます。</li>
		</ol>
		</li>
		<li>無菌技術を使用して、トリプシンのボリュームに MSCGM の同等のボリュームを含む 50 mL の円錐管に培養器から転送 hWJCs はセルを分離するに使用されます。</li>
		<li>常温 500 × g で遠心分離によって、hWJCs をペレット (~ 27 ° C) 5 分。</li>
		<li>上清を吸引し、500,000 細胞/ml の濃度で MSCGM で hWJCs を再懸濁します。</li>
	</ol>
	</li>
	<li>蝸牛を灌流します。
	<ol>
		<li>転送ピペットを使用して新しい 24 ウェル プレートに蝸牛を転送します。</li>
		<li>任意の蝸牛が乾燥するを防ぐために各の蝸牛に MSCGM のドロップを追加します。</li>
		<li>繊細な楕円形や円形の窓を向いているように超微細鉗子を使用して蝸牛に合わせます。 注: は、際は蝸牛が簡単に破損している各蝸牛を直接取り扱いに細心の注意を取る。</li>
		<li>接続チューブ、滅菌 28.5 ゲージ インスリン注射器の使用を再停止される 0.2 mL hWJCs (100,000 細胞) を描画します。</li>
		<li>滅菌微細鉗子を使用して、楕円形の窓にチューブを置きます。</li>
		<li>繊細かつゆっくり 0.2 ml ポリエチレン チューブに接続されている 28.5 ゲージ インスリン注射器を使用して再懸濁の hWJCs の蝸牛を灌流 (外径: 0.61 mm、内径: 0.28 mm) で約 5 分。</li>
		<li>後、0.8 mL 37 ° C の追加も 1 mL に井戸の総容積をもたらすに蝸牛を含む、MSCGM で加温します。</li>
		<li>各蝸牛、新鮮な注射器を使用して、各時間をチューブの 4.3.3-4.3.7 の手順を繰り返します。</li>
		<li>5% CO2、37 ° C 細胞文化のインキュベーターで灌流の蝸牛を置き、週 3 回のメディアを変更します。</li>
	</ol>
	</li>
</ol>
5. 蝸牛収穫と保存
注: 蝸牛培養があります、任意の時点で収穫まで 30 日後灌流。
<ol>
	<li>蝸牛を保持するには、ロッキング プラットフォーム上の 4 ° C で一晩 1x PBS で 4% パラホルムアルデヒドで修正します。</li>
	<li>固定後、3 回各 5 分の 1 × PBS と蝸牛を洗います。</li>
	<li>徐々 に脱水エタノールと蝸牛とのパラフィン包埋する前に脂肪族炭化水素溶剤でクリアします。</li>
	<li>セクションは、ミクロトームを使用して 10 μ m の厚さにサンプル、ガラス顕微鏡スライドをマウントします。 注: サンプルはする準備が整いました組織学的や免疫組織化学的処理の標準プロトコルを使用して。</li>
</ol>

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B. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Decellularized+ear+tissues+as+scaffolds+for+stem+cell+differentiation.">Decellularized ear tissues as scaffolds for stem cell differentiation.</a> J Assoc Res Otolaryngol. 14 (1), 3-15 (2013).</li><li>Davies, D., Holley, M. C. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Differential+expression+of+alpha+3+and+alpha+6+integrins+in+the+developing+mouse+inner+ear.">Differential expression of alpha 3 and alpha 6 integrins in the developing mouse inner ear.</a> J Comp Neurol. 445 (2), 122-132 (2002).</li><li>Gerchman, E., Hilfer, S. R., Brown, J. W. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Involvement+of+extracellular+matrix+in+the+formation+of+the+inner+ear.">Involvement of extracellular matrix in the formation of the inner ear.</a> Dev Dyn. 202 (4), 421-432 (1995).</li><li>Goodyear, R. J., Richardson, G. P. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Extracellular+matrices+associated+with+the+apical+surfaces+of+sensory+epithelia+in+the+inner+ear:+molecular+and+structural+diversity.">Extracellular matrices associated with the apical surfaces of sensory epithelia in the inner ear: molecular and structural diversity.</a> J Neurobiol. 53 (2), 212-227 (2002).</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=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>

著者が明らかに何もありません。

我々 は正常に蝸牛の複雑な三次元組織足場としての使用を可能にする decellularization プロセスを介して蝸牛からネイティブ蝸牛細胞を削除できることを実証しました。サンティら。15 decellularizing の蝸牛の最初のメソッドを開発し、光シート顕微鏡23の助けを借りてを通じて多くの人工内耳構造のボリュームを正確に推定しています。このような初期?...

蝸牛は、側頭骨は、複雑なスパイラル構造です。外側の骨迷路と、同心の内側の膜迷路1から成り立っています。膜迷路は、流体の 3 つのスペースで構成されています: Scala スカラー座 vestibuli メディア、およびスカラー座索1。Scala メディア住宅感覚の上皮細胞の種類の多数で構成されているが、感覚有毛細胞 (HC)、神経インパルス2音波の機械的エネルギーを変換、特に重要です。音響外傷3,4、5薬6疾患7,8、および高齢化9への暴露は、すべての HC の死を介して聴覚機能障害で起因できます。哺乳類の毛細胞損失は傷害10後再生することができます鳥の HCs とは異なり、永続的なです。
様々 な現代的な研究努力を失われた HCs を復元する特定の実験的アプローチは異なるが求めています。感覚上皮における遺伝子発現の操作と体の外分化した幹細胞の注入、支配的なアプローチにおいて、蝸牛オルガノイドの幹細胞の分化を誘導方法がされているが11,12,13を試みた。それぞれのアプローチは、直接幹細胞や幹細胞の発達の手がかりに依存ただし、2 番目は共有で、潜在的に重要な要素はそれ自体蝸牛の ECM14,15。
ECM だけでなく細胞と細胞の接着、増殖、生存、および移行の面を含んでいますが、また HCs やらせん神経節報15,16 の開発に重要な役割を果たしている組織の物理的なサポートを提供します ,17。自然発生する ECM は、細胞の表現型決定および/または細胞の接着・増殖・生存の18のガイドが誘導信号を提供します。その結果、脱蝸牛培養 hWJCs との組み合わせでの使用は、ECM と HC の再生の役割を探検するユニークな機会を提供しています。HWJCs は、間葉系幹細胞の19のように振る舞うひと臍帯由来容易に利用できる、非論争の携帯タイプです。HWJCs は、麻痺細胞系統20,21を区別する能力を示しています。したがって、現在のプロトコル詳細分離、decellularization、および内耳組織工学用 hWJCs と C57BL マウスの死骸から蝸牛の血流を示します。

<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

哺乳類のヒアリング不足を再生する能力を促進する機械刺激受容の有毛細胞では、難聴の治療法を限られています。現在の再生医療戦略は、移植幹細胞や内耳難聴を修正する損傷した幹細胞の交換を奨励するためにサポート細胞を周囲の遺伝子操作に焦点を当てています。しかし、細胞外マトリックス (ECM) は誘導と有毛細胞の機能を維持する重要な役割を果たすこと、十分に調査されていません。人工内耳を使用して成体幹細胞を成長する足場として ECM 方法、組成と細胞外の環境のアーキテクチャにエイズ聴覚機能の維持の細胞に独特な洞察力があります。分離し、成体幹細胞灌流を受け入れる足場として使用するマウスからの蝸牛を decellularizing する方法をご紹介します。現在のプロトコルの蝸牛は安楽死させたマウス、脱と脱灰から分離されます。その後、臍帯から隔離された人間の Wharton のゼリー細胞 (hWJCs) は、各蝸牛に慎重に灌。バイオリアクター, として、蝸牛を用い、解析の処理を受ける前に 30 日間培養します。脱の蝸牛は個人の細胞外構造を保持が、細胞の存在や DNA の顕著なフラグメントを明らかにしなかった。蝸牛に灌流セルは蝸牛の内外面のほとんどを侵略し、30 日間の期間にわたって問題なく育った。したがって、どのように人工内耳の ECM の影響細胞発達や行動を研究する現在のメソッドを使用することができます。

ここに提示されたメソッドを使用すると、蝸牛の成功した decellularization は、4' を介して DNA の存在有無を調べることで評価された 6-diamidino-2-phenylindole (DAPI) 染色します。Decellularized 蝸牛内 DNA が識別されなかった場合、蝸牛を完全に decellularized と考えられていた。Decellularization または脱灰を受けていない以前の実験からネイティブ蝸牛は、構造と伝統的 C57BL マウス蝸?...

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

内耳組織工学用マウス蝸牛を Decellularizing するためのプロトコル

このプロトコルの目的は、decellularize し、めしマウス蝸牛組織工学用の足場として利用する効果的な方法を示すことです。

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 ビュー.  University of Kansas Medical Center. このプロトコルの目的は、decellularize し、めしマウス蝸牛組織工学用の足場として利用する効果的な方法を示すことです。

ビデオ: 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.

心筋組織工学のための生物由来注射用材料の作製

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

生物工学

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.

肺のエンジニアリングのための手順

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.

動脈組織工学におけるエラストマーのPGSの足場

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.

心臓組織工学のためのフィブリンゲルにおける心筋細胞のカプセル化

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.

生体力学的解析と組織工学のための細胞由来の組織のリングを製作する監督セルラーセルフアセンブリ

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.

組織工学のためのエレクトロ繊維のポスト処理

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.

心臓組織工学のためのキャピラリーフォースリソグラフィ

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.

トライリーフレットポリマーバルブの相対的な流体力学的評価のためのプロトコル

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.

組織工学のための二軸メカニカルローディングバイオリアクターの設計

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.

エンジニアリング3D細胞化血管組織再生のためのコラーゲンゲル

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