D. H. Kim thanks the Department of Bioengineering at the University of Washington for the new faculty startup fund. D. H. Kim is also supported by the Perkins Coie Award for Discovery, the Wallace H. Coulter Foundation Translational Research Partnership Award, the Washington State Life Science Discovery Fund, and the American Heart Association Scientist Development Grant (13SDG14560076). J. Macadangdang and A. Jiao thank the support from the NIH Bioengineering Cardiovascular Training Grant Fellowship. &#160;Additional support for this work comes from the National Institutes of Health (NIH) grant R01HL111197 to M. Regnier.

特に断りのない限り、全ての手順を室温（〜23℃）で行われる。 シリコンマスターの1。作製<ol><li> O 2 / N 2ガ ​​ス下で100％エタノール、キシレン、ドライクリーンなシリコンウェーハ。 </li><li> 0.3〜0.5ミクロン厚のフィルムを製造するために2000〜4000回転の回転速度でスピンコーターでシリコンウエハを置く。 </li><li>フォトリソグラフィシステムを使用して、正しい寸法を有するパターンをフォトレジスト膜</li><li>完全にフォトレジスト現像液の適切な量のパターン化されたフォトレジストでコーティングされたシリコンウエハを浸す。 </li><li>脱イオン水で現像されたフォトレジストでコーティングされたシリコンウエハーを洗浄します。 </li><li>ほぼ垂直な側壁を有するサブミクロンスケールのリッジのアレイを形成するために、ディープ反応性イオンエッチング装置を用いて露出したシリコンをエッチングする。 </li><li>プラズマアッシャーシステムにおいてシリコンウェハを配置することによって、残りのフォトレジストを除去する。 </li><li> Wシリコンをカットその後のレプリカ成形に適したサイズのシリコンマスターズへのダイヤモンドが先端に付いたカッターでafers。 </li></ol>シリコン·マスターからのPUA型の2。作製注：ナノファブリケーションのためのシリコンマスターに追加されるボリュームは、ポリウレタンアクリレート（PUA）溶液の粘度ならびに複製されるナノパターンマスタの面積に応じて変化する。 <ol><li> 100％エタノール又はO 2 / N 2ガ ​​ス下、キシレン、ドライクリーンなシリコンマスタ面。 </li><li>シャーレにシ​​リコンマスターパターン側を上に置きます。 </li><li> 2センチメートル×2センチメートル表面パターン、パターン面への不要と思われるアプリケーションのピペット40μlのシリコンマスターのため。 </li><li> 4センチメートル注し、PUA上のX 4センチメートル透明なポリエステル（PET）フィルムのシートを配置します。 </li><li>ダウンPETシートを押して、そのように（カードなど）ローラーまたはフラットエッジの表面を用いたパターン面を横切るシートの下に不要と思われるアプリケーションを広める全体パターンは、PUAプレポリマーで覆われている。 </li><li> 50秒間約10cm 20ワット（115 V）のUV光（λ= 365 nm）を下回るシリコンマスター、プレポリマー、およびPETを配置します。効果的にするために、UV光の波長は310〜400nmの間の任意の場所とすることができる。光の強度は、基板表面の10〜15ミリワット/ cm 2である。 </li><li>硬化後、ピンセットでゆっくりとPETフィルムを除去します。 PUAは、シリコンマスターナノパターンの負のPETフィルムに添付してください。 </li><li>使用前に少なくとも12時間のUVで硬化PUA / PETナノパターン。露出オーバーは問題ではありません。 </li><li>シリコンマスターをクリーニングするには、不要と思われるアプリケーションを追加することなく、マスターの上にPETの別のフィルムを配置し、50秒間UV光（λ= 365 nm）に公開し、PETフィルムを除去します。これは、任意の未反応モノマーを除去します。 </li><li> O 2 / N 2ガ ​​ス下で100％エタノール、キシレン、乾燥してシリコンマスターを洗浄します。 </li></ol>図3（a）。ナノパターニングポリウレタンポリマー</P&gt; <ol><li> 10分間オゾン処理チャンバ内に配置することにより、直径25mmの円形ガラススライドを調製する。 </li><li>容易な取扱いのための小さなPDMSブロック上に、オゾン処理したガラススライドを置きます。 </li><li>ガラススライドに絵筆で表面接着促進剤の薄層を適用する。 30分間風乾燥したガラススライド。 </li><li>プリンタ紙にガラススライドを置きます。 </li><li>ドロップは、スライドガラスの中央部にポリウレタンの10μL（PU）プレポリマー（NOA 76）を分注する。気泡が添加後存在しないことを確認してください。 </li><li>場所のPUAモールド、パターンは、スライドガラス上に、下向きにセットしてください。 PUAモールドに沿ってゴム製のシリンダーローラーを転がすことにより、スライドガラスの表面全体に均一にPUを分散させる。プリンタ用紙は、ポリマーのオーバーフローを吸収する。 </li><li>ワットUVランプ。 PUの重合時間は、UV光源のパワーに依存している。 </li><li> UV光源からサンプルを取り出して、慎重にPUコーティングされたガラススライドからPUA金型をはがします。 PUの重合は、Cと考えられているompleteとききれいにサンプルからPUAモールドの皮とPUガラススライドは虹色の外観を有する。 </li><li>限りの月のようにストレージのためにデシケーターに仕上がったサンプルを置きます。 </li></ol>図3（b）。ナノパターニングポリ（ラクチド - コ - グリコリド）ヒドロゲル<ol><li>激しく10:1の比でシリコーンエラストマーベース及びシリコーンエラストマー硬化剤を混合することにより、平坦なPDMSモールドを作成する。 </li><li> PDMS前駆体は皿（ すなわち、PDMSを厚さ5mmになるように）の縁5mmまで到達するようにペトリ皿に混合PDMS前駆体溶液を注ぐ。 </li><li>脱ガスに1時間デシケーター中でペトリ皿とPDMS前駆体を配置します。 </li><li>治すために、少なくとも2時間、65℃のオーブンにペトリ皿とPDMS前駆体を移動します。 </li><li> PDMSを硬化させた後は3センチ×3センチメートルの正方形の部分に平らなPDMSをカットするかみそりを使用しています。これらは、パターニングプロセスの後半で使用されるフ​​ラットなPDMSモールドとなります。 </li><li> PLによるクリーンな直径25mmの円形のガラススライド水中超音波処理装置での30分間イソプロピルアルコールでacing。 </li><li> O 2 / N 2ガ ​​ス下で乾燥した清浄なガラススライド。 </li><li>液滴をガラススライド上にPLGA溶液100μl（クロロホルム中/ vのPLGA 15重量％）を分注する。 </li><li>溶剤を吸収し、平らなPLGA表面を得るに分注し、PLGAの上に平らなPDMSモールドを配置します。 5分間PDMSの上に200グラムの重量を配置することによって、光の圧力（〜10キロパスカル）を適用します。 </li><li>ゆっくりと平坦なPDMSモールドを剥離し、残留溶媒を除去し、PLGAとカバー​​ガラスとの密着性を高めるために5分間予熱したプレート（120°C）でカバーガラスを配置する。 </li><li>フラットPLGAの上のNP PUAモールドを配置し、15分間の加熱プレート上にある間、PUAモールドの上に1kgの重りを配置することによって、一定の圧力（〜100キロパスカル）と熱（120℃）を適用する。 </li><li> PLGAカバースライドから重量を除去し、基質を室温まで冷却する。基層がされるまで、PUAモールドを削除しないでください冷却した。 </li><li>基質が十分に冷却した後、慎重にNPのPLGA基層を明らかにし、NP PUAモールドをはがす。 NPのPLGA基層は、虹色の外観を持っている必要があります。 </li><li>限りの月のようにストレージのためにデシケーターに仕上がったサンプルを置きます。 </li></ol> 4。細胞播種および培養注：このプロトコルは、新生児ラット心室筋細胞（NRVMs）およびH7ヒト胚性幹細胞由来の心筋細胞（hESCの-CMS）が、他の細胞源の文化を説明して使用することができる。 <ol><li> 35mmの組織培養ポリスチレン皿にANFSカバースリップ（PUまたはPLGA）を取り付けます。ピペットノーランド光学接着剤（NOA83H）20μlの皿の底に、ゆっくりとNOAの上にANFSカバースリップを配置。接着剤が広がって全体のカバーガラス底をカバーすることができます。 10分間UVに料理を暴露することによりNOAを治す。 </li><li>二回、5分間、70％エタノール水溶液2mlですすぐことによりANFSを滅菌する。 Eを削除吸引によりthanol。 ANF​​Sは生物学的安全キャビネット内のUV殺菌ランプ（λ= 200から290ナノメートル）の下で〜1時間、完全に空気乾燥させます。 </li><li>細胞接着を一晩フィブロネクチンでANFSを被覆することによって強化される。 5μg/ mlのDI水にフィブロネクチンを希釈する。皿にフィブロネクチン溶液をピペット2ミリリットル。 37℃のインキュベーター中に置き、5％CO 2で一晩（少なくとも6時間）。 </li><li>以前プロトコル22に従ってNRVMs、ヒトES細胞-CMS、または関心のある他の心臓細胞を得る。 </li><li>細胞をペレットに3分間1,000 rpmで遠心分離細胞サンプル。 </li><li>慎重に吸引して上清を除去。ペレットを乱さないように注意してください。 </li><li> 4.6×10 6細胞/ mlの濃度になるように適切な培養培地中で細胞を再懸濁。 </li><li>慎重に滅菌ANFS上に細胞懸濁液200μlをピペットで。細胞懸濁液は、カバーガラス上に残っていることを確認してください。 </li><li> 37℃、5％のインキュベーター中で細胞を置き細胞はANFSにアタッチできるようにする4時間のCO 2。 </li><li>同じ条件で培養器で細胞を皿に代わる暖かい培養培地の2ミリリットルを追加します。 </li><li> 24時間後、培地を除去し、余分な細胞を除去するために二度のDPBS 2ミリリットルで洗浄する。 </li><li>同じ条件で培養器で細胞を皿に代わる暖かい培養培地の2ミリリットルを追加します。文化細胞が密集するまで。一日おきにメディアを交換してください。 </li></ol>

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W., 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=Biological+length+scale+topography+enhances+cell-substratum+adhesion+of+human+corneal+epithelial+cells.">Biological length scale topography enhances cell-substratum adhesion of human corneal epithelial cells.</a> J Cell Sci. 117, 3153-3164 (2007).</li><li>Cavalcanti-Adam, E. A., 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=Cell+Spreading+and+Focal+Adhesion+Dynamics+Are+Regulated+by+Spacing+of+Integrin+Ligands.">Cell Spreading and Focal Adhesion Dynamics Are Regulated by Spacing of Integrin Ligands.</a> Biophysical Journal. 92, 2964-2974 (2007).</li><li>Koo, L. Y., Irvine, D. J., Mayes, A. M., Lauffenburger, D. A., Griffith, L. 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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=The+control+of+human+mesenchymal+cell+differentiation+using+nanoscale+symmetry+and+disorder.">The control of human mesenchymal cell differentiation using nanoscale symmetry and disorder.</a> Nat Mater. 6, 997-1003 (2007).</li><li>Park, J., Bauer, S., Mark, v. o. n. . d. e. r. . 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Authors have nothing to disclose.

機能的に成熟した心臓組織は、 生体内および心臓組織工学のインビトロ適用の両方のために欠けている。ここで説明するのCFLナノ加工法は、細胞の位置合わせを実現し、システムの拡張性に起因する巨視的組織機能に影響を与えるための堅牢な技術である。大きな面積を容易にパターニングし、細胞培養のために使用することができる。巨視的セルラアラインメントは生...

心血管疾患は、世界における罹患率および死亡率の主要な原因であり、すでに緊張した世界的な保健システム1,17上の重い社会経済的な負担を提示する。 （1）in vitroでの薬物スクリーニングまたは疾患のモデリングのための心臓の高忠実度モデルを作成する虚血性疾患又は心筋症又は（2）の後に損傷した心筋を再生する：心臓組織工学は、2つの異なる目標を有している。 心臓は体に血液を供給するために常に働かなければならない複雑な器官である。心筋細胞および支持組織の密集した層構造は、心臓壁18,19を通して螺旋パターンで配置されている。心臓はまた、電気機械的に効率的に本体21に血液を排出する高度に配位様式20に結合される。自然の複雑なデザインを確実vitroでで要約することができる前に、いくつかの主要なハードルは、しかし、対処されていない。ロバストな心筋細胞分化方法22開発され続けているが、まず、HPSC-CMはまだかなり未成熟表現型を示す。彼らの電気機械特性および形態は、最も密接に胎児のレベルを23に一致する。従来の培養条件で維持さときに、第2、幹細胞由来の心筋細胞および一次両方は、天然、組織様構造に組み立てることができない。むしろ、細胞がランダムに配向となり、大人の心筋24のバンド棒状の外観を示さない。 細胞が相互作用する細胞外マトリックス（ECM）環境では、多数の細胞プロセス11,13,25において重要な役割を果たしている。 ECMは、細胞を有意に6,26の構造及び機能に影響を与える複雑な、明確に定義された分子及び地形手がかりからなる。心の中では、携帯の位置合わせは、密接に、基礎となるナノメートルスケールのECM繊維2に従います。これらnanotopographの影響細胞および組織の機能上のiCalの手がかりは、しかし、完全には理解遠くからである。ナノメートルスケール細胞生体材料の相互作用の予備的研究は、成長31、接着28〜30、27細胞シグナル伝達のためにサブミクロン地形手がかりの潜在的な重要性と影響力を示し、分化32,33。しかし、再現性のあるスケーラブルなnanofabricated基板を開発することが困難であるため、このような研究は、 生体内のECM環境での複合体のマルチスケール細胞効果を再現することができませんでした。このプロトコルでは、天然の心臓ECM繊維の配向を模倣する細胞培養足場を製造する簡単で費用効果の高いナノファブリケーション技術は、心筋細胞、生体材料の相互作用の新規な調査の広い範囲を可能にする、記載されている。心筋細胞は、ナノスケールのECM環境とどのように相互作用するか理解することは、より密接に模倣天然組織の機能に細胞挙動を制御する能力のために可能性がありますTiONから。さらに、細胞単層は、3D構造に比べて簡略化された実験系であるが、それでも洞察に満ちた調査と機能的スクリーニング2,34-36ための複雑な多細胞挙動を示す。最後に、このような足場は、再生目的のために心臓37内に埋め込 ​​まれた際の細胞移植片の機能を改善するために使用することができる。

<table border="0" cellpadding="0" cellspacing="0" width="352">
	<colgroup>
		<col span="4" style="width:66pt" width="88" />
	</colgroup>
	<tbody>
		<tr height="43" style="height:32.25pt">
			<td class="xl63" height="43" style="height:32.25pt;width:66pt" width="88">Name</td>
			<td class="xl63" style="width:66pt" width="88">Company</td>
			<td class="xl63" style="width:66pt" width="88">Catalogue number</td>
			<td class="xl64" style="width:66pt" width="88">Comments (optional)</td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Fibronectin</td>
			<td class="xl65" style="width:66pt" width="88">BD Biosciences</td>
			<td class="xl65" style="width:66pt" width="88">354008</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="64" style="height:48.0pt">
			<td class="xl65" height="64" style="height:48.0pt;width:66pt" width="88">NOA 76</td>
			<td class="xl65" style="width:66pt" width="88">Norland Products, Inc.</td>
			<td class="xl65" style="width:66pt" width="88">7606B</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="106" style="height:79.5pt">
			<td class="xl65" height="106" style="height:79.5pt;width:66pt" width="88">Surface Adhesion Promotor (Glass Primer)</td>
			<td class="xl65" style="width:66pt" width="88">Minuta Tech</td>
			<td class="xl65" style="width:66pt" width="88"></td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">PUA</td>
			<td class="xl65" style="width:66pt" width="88">Minuta Tech</td>
			<td class="xl65" style="width:66pt" width="88">MINS-311RM</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Soft Rubber Roller</td>
			<td class="xl65" style="width:66pt" width="88">Speedball</td>
			<td class="xl65" style="width:66pt" width="88"></td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="64" style="height:48.0pt">
			<td class="xl65" height="64" style="height:48.0pt;width:66pt" width="88">Silicon Wafers</td>
			<td class="xl65" style="width:66pt" width="88">NOVA Electronic Materials</td>
			<td class="xl65" style="width:66pt" width="88">FA01-9900</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="22" style="height:16.5pt">
			<td class="xl65" height="22" style="height:16.5pt;width:66pt" width="88">Photoresist</td>
			<td class="xl65" style="width:66pt" width="88">Shipley</td>
			<td class="xl65" style="width:66pt" width="88">SPRT510</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Photoresist Developer</td>
			<td class="xl65" style="width:66pt" width="88">Shipley</td>
			<td class="xl65" style="width:66pt" width="88">MF320</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="85" style="height:63.75pt">
			<td class="xl65" height="85" style="height:63.75pt;width:66pt" width="88">Electron-Beam Lithography System</td>
			<td class="xl65" style="width:66pt" width="88">JEOL</td>
			<td class="xl65" style="width:66pt" width="88">JBX-9300FS</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="64" style="height:48.0pt">
			<td class="xl65" height="64" style="height:48.0pt;width:66pt" width="88">Etching System</td>
			<td class="xl65" style="width:66pt" width="88">Surface Technology Systems</td>
			<td class="xl65" style="width:66pt" width="88">NP10 8UJ</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="64" style="height:48.0pt">
			<td class="xl65" height="64" style="height:48.0pt;width:66pt" width="88">Plasma Asher System</td>
			<td class="xl65" style="width:66pt" width="88">BMR Technology Co.</td>
			<td class="xl65" style="width:66pt" width="88">DSF-200</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Ozone Cure System</td>
			<td class="xl65" style="width:66pt" width="88">Minuta Tech</td>
			<td class="xl65" style="width:66pt" width="88">MT-UV-O- 08</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Fusion Cure System</td>
			<td class="xl65" style="width:66pt" width="88">Minuta Tech</td>
			<td class="xl65" style="width:66pt" width="88">MT-UV-A 11</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="64" style="height:48.0pt">
			<td class="xl65" height="64" style="height:48.0pt;width:66pt" width="88">NOA 83H</td>
			<td class="xl65" style="width:66pt" width="88">Norland Products, Inc.</td>
			<td class="xl65" style="width:66pt" width="88">8301</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Spin Coater</td>
			<td class="xl65" style="width:66pt" width="88">Laurel Technology</td>
			<td class="xl65" style="width:66pt" width="88">WS-400-6NPP</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Skyrol PET Film</td>
			<td class="xl65" style="width:66pt" width="88">SKC Co., Ltd.</td>
			<td class="xl65" style="width:66pt" width="88">23038-59-9</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">25mm Glass Slides</td>
			<td class="xl65" style="width:66pt" width="88">Corning</td>
			<td class="xl65" style="width:66pt" width="88">2948</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="85" style="height:63.75pt">
			<td class="xl65" height="85" style="height:63.75pt;width:66pt" width="88">Sylgard 184 Silicone Elastomer Kit</td>
			<td class="xl65" style="width:66pt" width="88">Dow Corning</td>
			<td class="xl67" style="width:66pt" width="88">6/5/2553</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="64" style="height:48.0pt">
			<td class="xl65" height="64" style="height:48.0pt;width:66pt" width="88">Poly(D,L-lactide-co-glycolide)</td>
			<td class="xl65" style="width:66pt" width="88">Sigma-Aldrich</td>
			<td class="xl65" style="width:66pt" width="88">P2191-1G</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Chloroform</td>
			<td class="xl65" style="width:66pt" width="88">Sigma-Aldrich</td>
			<td class="xl65" style="width:66pt" width="88">372978-1L</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">500g Weights</td>
			<td class="xl65" style="width:66pt" width="88">Global Insustrial</td>
			<td class="xl65" style="width:66pt" width="88">T9FB503120</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="43" style="height:32.25pt">
			<td class="xl65" height="43" style="height:32.25pt;width:66pt" width="88">Isopropyl Alcohol</td>
			<td class="xl65" style="width:66pt" width="88">EMD Millipore</td>
			<td class="xl65" style="width:66pt" width="88">PX1835-2</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="22" style="height:16.5pt">
			<td class="xl65" height="22" style="height:16.5pt;width:66pt" width="88">Hot Plate</td>
			<td class="xl65" style="width:66pt" width="88">Corning</td>
			<td class="xl65" style="width:66pt" width="88">PC-420D</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
		<tr height="22" style="height:16.5pt">
			<td class="xl65" height="22" style="height:16.5pt;width:66pt" width="88">Sonicator</td>
			<td class="xl65" style="width:66pt" width="88">Branson</td>
			<td class="xl65" style="width:66pt" width="88">B2510MTH</td>
			<td class="xl66" style="width:66pt" width="88"></td>
		</tr>
	</tbody>
</table>

capillary force lithography for cardiac tissue engineering

心血管疾患は世界的に1死亡の主要な原因である。心臓組織工学は、心臓再生だけでなく、in vitroスクリーニングアッセイのための機能的な組織を開発する目的で画期的な医療発見を実現するために多くの可能性を秘めています。しかしながら、心臓組織の高忠実度のモデルを作成することが困難であることが証明されている。心臓の細胞外マトリックス（ECM）は、マイクロナノメートルスケールの2に至るまでの生化学的および生体力学的信号の両方からなる複合構造体である。ローカル機械的負荷条件及び細胞-ECM相互作用は、最近、心臓組織工学3-5の重要なコンポーネントとして認識されている。 心臓ECMの大部分は有意に組織構造および電気機械結合に影響を与える2ナノスケールの直径を有する整列されたコラーゲン繊維から構成されている。残念ながら、いくつかのメソッドは、HAVEナノメートルスケールまでECM繊維の組織を模倣することができた。ナノファブリケーション技術の最近の進歩は、しかし、心の6-9でのECMのin vivoでの構造と基板の剛性の手がかりを模倣する、スケーラブルな足場の設計および製造を可能にした。 ここでは、の開発に存在する二つの再現可能な、費用対効果、および生体適合性ポリマーであるポリ（ラクチド-コ-グリコリド）（PLGA）8およびポリウレタン（PU）系ポリマーを用いて心臓細胞の機能的な位置合わせのためのスケーラブルなナノパターニングプロセス。これらの異方nanofabricated基層（ANFS）はよく組織化、整列の組織の根本的なECMを模倣し、細胞の形態と機能10から14にナノトポの役割を調査するために使用できます。 鋳型としてナノパターン（NP）シリコンマスターを使用して、ポリウレタンアクリレート（PUA）モールドが作製される。このPUAモールドは、その後、PAに使用されているそれぞれ、UV支援または溶媒媒介毛細管力リソグラフィー（CFL）、15,16を経由して、PUまたはPLGAヒドロゲルをttern。簡潔には、PUまたはPLGAプレポリマーは、ドロップガラスカバースリップ上に分注し、PUAモールドを上に置くである。 UVアシストCFLの場合、PUは、次いで硬化させるためのUV照射（λ= 250〜400 nm）に露光される。溶媒媒介性CFLについては、PLGAを、熱（120℃）および圧力（100kPaで）を用いてエンボス加工される。硬化後、PUAモールドを細胞培養のためのANFS残して、剥離される。例えば、新生仔ラット心室筋細胞、ならびにヒト多能性幹細胞由来の心筋細胞のような初代細胞は、ANFS 2上に維持することができる。

図1は、2つの製造方法の製造工程の概略図である。によりナノスケールの地形に起因する光の回折に、ナノパターニングはANFSに虹色の表面をもたらすはずである。 図2は、800 nmの稜線と溝とよくパターン化された25ミリメートルのNP-PUカバースリップ（図2A）に、この虹色の表面を示している幅（図2B）。 ANF​​Sの虹色の外観は、畝と溝の幅?...

Watch this Scientific Journal Video about Capillary Force Lithography for Cardiac Tissue Engineering at JoVE.com

Capillary Force Lithography for Cardiac Tissue Engineering

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 ...

capillary-force-lithography-for-cardiac-tissue-engineering

Research

JoVE Journal

Bioengineering

12.3K ビュー.  University of Washington. 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.

ビデオ: Capillary Force Lithography for Cardiac 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

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 ...

Plasma Lithography Surface Patterning for Creation of Cell Networks

Systematic manipulation of a cell microenvironment with micro- and nanoscale resolution is often required for deciphering various cellular and molecular phenomena. To address this requirement, we have developed a plasma lithography technique to manipulate the cellular microenvironment by creating a patterned surface with feature sizes ranging from 100 nm to millimeters. The goal of this technique is to be able to study, in a controlled way, the behaviors of individual cells as well as groups of cells and their interactions.
This plasma lithography method is based on selective modification of the surface chemistry on a substrate by means of shielding the contact of low-temperature plasma with a physical mold. This selective shielding leaves a chemical pattern which can guide cell attachment and movement. This pattern, or surface template, can then be used to create networks of cells whose structure can mimic that found in nature and produces a controllable environment for experimental investigations. The technique is well suited to studying biological phenomenon as it produces stable surface patterns on transparent polymeric substrates in a biocompatible manner. The surface patterns last for weeks to months and can thus guide interaction with cells for long time periods which facilitates the study of long-term cellular processes, such as differentiation and adaption. The modification to the surface is primarily chemical in nature and thus does not introduce topographical or physical interference for interpretation of results. It also does not involve any harsh or toxic substances to achieve patterning and is compatible for tissue culture. Furthermore, it can be applied to modify various types of polymeric substrates, which due to the ability to tune their properties are ideal for and are widely used in biological applications. The resolution achievable is also beneficial, as isolation of specific processes such as migration, adhesion, or binding allows for discrete, clear observations at the single to multicell level.
This method has been employed to form diverse networks of different cell types for investigations involving migration, signaling, tissue formation, and the behavior and interactions of neurons arraigned in a network.

A versatile plasma lithography technique has been developed to generate stable surface patterns for guiding cellular attachment. This technique can be applied to create cell networks including those that mimic natural tissues and has been used for studying several, distinct cell types.

細胞ネットワークの作成のためのプラズマのリソグラフィ表面パターニング

Systematic manipulation of a cell microenvironment with micro- and nanoscale resolution is often required for deciphering various cellular and molecular ...

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 ...

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 ...

Tissue Engineering of a Human 3D in vitro Tumor Test System

Cancer is one of the leading causes of death worldwide. Current therapeutic strategies are predominantly developed in 2D culture systems, which inadequately reflect physiological conditions in vivo. Biological 3D matrices provide cells an environment in which cells can self-organize, allowing the study of tissue organization and cell differentiation. Such scaffolds can be seeded with a mixture of different cell types to study direct 3D cell-cell-interactions. To mimic the 3D complexity of cancer tumors, our group has developed a 3D in vitro tumor test system.	
	Our 3D tissue test system models the in vivo situation of malignant peripheral nerve sheath tumors (MPNSTs), which we established with our decellularized porcine jejunal segment derived biological vascularized scaffold (BioVaSc). In our model, we reseeded a modified BioVaSc matrix with primary fibroblasts, microvascular endothelial cells (mvECs) and the S462 tumor cell line. For static culture, the vascular structure of the BioVaSc is removed and the remaining scaffold is cut open on one side (Small Intestinal Submucosa SIS-Muc). The resulting matrix is then fixed between two metal rings (cell crowns).	
	Another option is to culture the cell-seeded SIS-Muc in a flow bioreactor system that exposes the cells to shear stress. Here, the bioreactor is connected to a peristaltic pump in a self-constructed incubator. A computer regulates the arterial oxygen and nutrient supply via parameters such as blood pressure, temperature, and flow rate. This setup allows for a dynamic culture with either pressure-regulated pulsatile or constant flow.	
	In this study, we could successfully establish both a static and dynamic 3D culture system for MPNSTs. The ability to model cancer tumors in a more natural 3D environment will enable the discovery, testing, and validation of future pharmaceuticals in a human-like model.

Tissue Engineering of a Human 3D in vitro Tumor Test System

Methods to create human 3D tumor tissues as test systems are described. These technologies are based on a decellularized Biological Vascularized Scaffold (BioVaSc), primary human cells and a tumor cell line, which can be cultured under static as well as under dynamic conditions in a flow bioreactor.

人間の3次元の組織工学<em&gt; in vitroで</em&gt;腫瘍テストシステム

Cancer is one of the leading causes of death worldwide.  Current therapeutic strategies are predominantly developed in 2D culture  systems, which ...

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. ...

Preparation of Thermoresponsive Nanostructured Surfaces for Tissue Engineering

Thermoresponsive poly(N-isopropylacrylamide) (PIPAAm)-immobilized surfaces for controlling cell adhesion and detachment were fabricated by the Langmuir-Schaefer method. Amphiphilic block copolymers composed of polystyrene and PIPAAm (St-IPAAms) were synthesized by reversible addition-fragmentation chain transfer (RAFT) radical polymerization. A chloroform solution of St-IPAAm molecules was gently dropped into a Langmuir-trough apparatus, and both barriers of the apparatus were moved horizontally to compress the film to regulate its density. Then, the St-IPAAm Langmuir film was horizontally transferred onto a hydrophobically modified glass substrate by a surface-fixed device. Atomic force microscopy images clearly revealed nanoscale sea-island structures on the surface. The strength, rate, and quality of cell adhesion and detachment on the prepared surface were modulated by changes in temperature across the lower critical solution temperature range of PIPAAm molecules. In addition, a two-dimensional cell structure (cell sheet) was successfully recovered on the optimized surfaces. These unique PIPAAm surfaces may be useful for controlling the strength of cell adhesion and detachment.

Nanoscaled sea-island surfaces composed of thermoresponsive block copolymers were fabricated by the Langmuir-Schaefer method for controlling spontaneous cell adhesion and detachment. Both the preparation of the surface and the adhesion and detachment of cells on the surface were visualized.

組織工学のための熱応答性ナノ構造表面の調製

Thermoresponsive poly(N-isopropylacrylamide) (PIPAAm)-immobilized surfaces for controlling cell adhesion and detachment were fabricated by the ...

Tissue Engineering by Intrinsic Vascularization in an In Vivo Tissue Engineering Chamber

In reconstructive surgery, there is a clinical need for an alternative to the current methods of autologous reconstruction which are complex, costly and trade one defect for another. Tissue engineering holds the promise to address this increasing demand. However, most tissue engineering strategies fail to generate stable and functional tissue substitutes because of poor vascularization. This paper focuses on an in vivo tissue engineering chamber model of intrinsic vascularization where a perfused artery and a vein either as an arteriovenous loop or a flow-through pedicle configuration is directed inside a protected hollow chamber. In this chamber-based system angiogenic sprouting occurs from the arteriovenous vessels and this system attracts ischemic and inflammatory driven endogenous cell migration which gradually fills the chamber space with fibro-vascular tissue. Exogenous cell/matrix implantation at the time of chamber construction enhances cell survival and determines specificity of the engineered tissues which develop. Our studies have shown that this chamber model can successfully generate different tissues such as fat, cardiac muscle, liver and others. However, modifications and refinements are required to ensure target tissue formation is consistent and reproducible. This article describes a standardized protocol for the fabrication of two different vascularized tissue engineering chamber models in vivo.

Tissue Engineering by Intrinsic Vascularization in an In Vivo Tissue Engineering Chamber

This is a guideline for constructing in vivo vascularized tissue using a microsurgical arteriovenous loop or a flow-through pedicle configuration inside a tissue engineering chamber. The vascularized tissues generated can be employed for organ regeneration and replacement of tissue defects, as well as for drug testing and disease modeling.

中に組み込み血管新生による組織工学<em&gt;インビボ</em&gt;ティッシュ・エンジニアリング商工会議所

In reconstructive surgery, there is a clinical need for an alternative to the current methods of autologous reconstruction which are complex, costly and ...

A Protocol for Decellularizing Mouse Cochleae for Inner Ear Tissue Engineering

In mammals, mechanosensory hair cells that facilitate hearing lack the ability to regenerate, which has limited treatments for hearing loss. Current regenerative medicine strategies have focused on transplanting stem cells or genetic manipulation of surrounding support cells in the inner ear to encourage replacement of damaged stem cells to correct hearing loss. Yet, the extracellular matrix (ECM) may play a vital role in inducing and maintaining function of hair cells, and has not been well investigated. Using the cochlear ECM as a scaffold to grow adult stem cells may provide unique insights into how the composition and architecture of the extracellular environment aids cells in sustaining hearing function. Here we present a method for isolating and decellularizing cochleae from mice to use as scaffolds accepting perfused adult stem cells. In the current protocol, cochleae are isolated from euthanized mice, decellularized, and decalcified. Afterward, human Wharton&#39;s jelly cells (hWJCs) that were isolated from the umbilical cord were carefully perfused into each cochlea. The cochleae were used as bioreactors, and cells were cultured for 30 days before undergoing processing for analysis. Decellularized cochleae retained identifiable extracellular structures, but did not reveal the presence of cells or noticeable fragments of DNA. Cells perfused into the cochlea invaded most of the interior and exterior of the cochlea and grew without incident over a duration of 30 days. Thus, the current method can be used to study how cochlear ECM affects cell development and behavior.

The goal of this protocol is to demonstrate an effective method to decellularize and decalcify mouse cochleae for utilization as scaffolds for tissue engineering applications.

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

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