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.

Alle Verfahren wurden bei Raumtemperatur (~ 23 ° C) durchgeführt, sofern nicht anders angegeben. 1. Herstellung von Silicon-Master <ol><li> Saubere Silizium-Wafer mit 100% Ethanol oder Xylol und trocken unter O 2 / N 2-Gas. </li><li> Zeigen Silizium-Wafer in Spin-Coater bei Drehzahlen von 2.000-4.000 Umdrehungen pro Minute, um eine 0,3-0,5 um dicken Film zu produzieren. </li><li> Strukturieren der Photoresistschicht mit den richtigen Dimensionen unter Verwendung eines Photolithographiesystem </li><li> Vollständig eintauchen, die strukturierte Photoresist-beschichtete Siliziumwafer in einem geeigneten Volumen von Photoresistentwicklerlösung. </li><li> Spülen Sie die entwickelte Photoresist beschichteten Silizium-Wafer mit VE-Wasser. </li><li> Arrays von Sub-Mikrometer-Maßstab Rippen mit fast vertikalen Seitenwänden, tiefes reaktives Ionenätzen der freiliegenden Silizium unter Verwendung eines Ätz-System zu bilden. </li><li> Entfernen Sie die verbleibende Photoresist, indem die Silizium-Wafer in einem Plasmaverascher-System. </li><li> Schneiden Sie die Silizium wAfers mit einem Diamantschneider in die entsprechend dimensioniert Silizium-Meister für die anschließende Replik Form. </li></ol> 2. Herstellung von PUA-Mold von Silicon Meister HINWEIS: Volumen auf Silizium-Master zur Nanofabrikation zugesetzt werden in Abhängigkeit von der Fläche der nano Master als auch repliziert werden, da die Viskosität des Polyurethanacrylat (PUA)-Lösung variieren. <ol><li> Saubere Silizium-Master-Oberfläche mit 100% Ethanol oder Xylol und trocken unter O 2 / N 2-Gas. </li><li> Zeigen Silizium-Master-Muster Seite nach oben in eine Petrischale. </li><li> Für eine Silizium-Master mit einer 2 cm x 2 cm Flächenmuster, Pipette 40 ul PUA auf die Musteroberfläche. </li><li> Legen Sie ein Blatt von 4 cm x 4 cm transparentem Polyester (PET) über das abgegebene PUA. </li><li> Drücken Sie auf die PET-Folie und die Ausbreitung der PUA unter dem Bogen über die Musterfläche mit einer Rolle oder Flachkantige Oberfläche (wie eine Karte), so dass der gesamteMuster durch die PUA-Prepolymer bedeckt. </li><li> Zeigen Silizium-Master, Prepolymer und PET etwa 10 cm unter einer 20 Watt (115 V) UV-Licht (λ = 365 nm) für 50 Sekunden. Um wirksam zu sein, kann das UV-Licht der Wellenlänge 310 bis 400 nm irgendwo zwischen sein. Die Intensität des Lichts 10-15 mW / cm 2 an der Oberfläche des Substrats. </li><li> Nach dem Aushärten entfernen PET-Folie mit einer Pinzette langsam. PUA sollte auf die PET-Folie mit einer negativen des Silizium-Master Nano befestigen. </li><li> Cure PUA / PET Nano unter UV für mindestens 12 Stunden vor der Verwendung. Überbelichtung ist nicht ein Problem. </li><li> So reinigen Silizium Meister, legen Sie eine weitere Folie aus PET auf dem Master ohne Zusatz von PUA und setzen UV-Licht (λ = 365 nm) für 50 sec und PET-Folie entfernen. Dadurch werden alle nicht umgesetzten Monomere zu entfernen. </li><li> Spülen Silizium-Master mit 100% Ethanol oder Xylol und trocken unter O 2 / N 2-Gas. </li></ol> 3a. Nanostrukturierung Polyurethan-Polymer </p&gt; <ol><li> Bereiten Durchmesser von 25 mm runden Glasobjektträger, indem sie in einer Ozonbehandlungskammer für 10 min. </li><li> Zeigen Ozon-behandelten Glasobjektträger auf kleine PDMS-Block für einfache Handhabung. </li><li> Eine dünne Schicht von Oberflächenhaftvermittler mit Pinsel auf Glasobjektträger. Air trockenen Glasobjektträger für 30 min. </li><li> Setzen Sie den Glasobjektträger auf einem Stück Druckerpapier. </li><li> Lassen Sie verzichten 10 ul von Polyurethan (PU) Pre-Polymer (NOA 76) bis zur Mitte des Glasträger. Stellen Sie sicher, dass keine Luftblasen vorhanden sind nach der Zugabe. </li><li> Platz PUA Form, Muster Gesicht nach unten, auf den Glasträger. Dispergieren des PU gleichmäßig über die Oberfläche der Glasplatte durch Walzen eines Walzengummizylinder entlang der PUA Form. Druckerpapier wird Polymer Lauf absorbieren. </li><li> Watt UV Lampe. Polymerisation Zeitpunkt der PU ist abhängig von Leistung von UV-Quelle. </li><li> Entfernen Sie die Probe von UV-Lichtquelle und vorsichtig schälen die PUA-Form aus PU-beschichteten Objektträger. Die Polymerisation des PU wird als complete wenn die PUA Formschalen sauber von der Probe und die PU-Objektträger aus Glas hat eine schillernde Erscheinung. </li><li> Platzieren abgeschlossen Proben Exsikkator zur Lagerung solange einen Monat. </li></ol> 3b. Nano Poly (lactid-co-glycolid)-Hydrogel <ol><li> Erstellen Sie einen Flach PDMS-Form durch kräftiges Mischen Silikonelastomer Basis-und Silikon-Elastomer Härter in einem Verhältnis von 10:1. </li><li> Gießen gemischten Vorläufer PDMS-Lösung in eine Petrischale, so dass der PDMS-Vorstufe erreicht 5 mm den Rand der Schale (also so, dass die PDMS ist 5 mm dick). </li><li> Zeigen Petrischale und PDMS-Vorläufer in einem Exsikkator für 1 h entgast. </li><li> Petrischale bewegen und PDMS-Vorläufers in einem 65 ° C Ofen für mindestens 2 Stunden aushärten. </li><li> Nach PDMS ausgehärtet ist, verwenden Sie einen Rasierer, um die flachen PDMS in 3 cm x 3 cm quadratische Abschnitte geschnitten. Diese werden die flachen Formen PDMS später in der Strukturierungsprozess verwendet wird. </li><li> Saubere 25 mm Durchmesser kreisförmige Glasobjektträger durch placing in Isopropylalkohol für 30 Minuten in einem Wasserbad-Beschallungs. </li><li> Trocken gereinigt Glasobjektträger unter O 2 / N 2-Gas. </li><li> Drop werden 100 ul PLGA-Lösung (15% w / v PLGA in Chloroform) auf Glasobjektträger. </li><li> Legen Sie einen flachen PDMS-Form auf der Oberseite der abgegebenen PLGA, um das Lösungsmittel zu absorbieren und zu erhalten, eine flache Oberfläche PLGA. Tragen Sie einen leichten Druck (~ 10 kPa), indem Sie einen 200 g Gewicht auf der PDMS für 5 min. </li><li> Langsam lösen sich flach PDMS-Form und legen Sie das Deckglas auf einer vorgewärmten Platte (120 ° C) für 5 min auf Restlösemittel zu entfernen und erhöhen die Haftung zwischen der PLGA und dem Deckglas. </li><li> Setzen Sie den NP PUA Form auf der Oberseite des Flach PLGA und gelten konstanten Druck (~ 100 kPa) und Hitze (120 ° C), indem Sie einen 1 kg-Gewicht auf der Oberseite des PUA-Form, während auf der Heizplatte für 15 min. </li><li> Entfernen Sie das Gewicht von PLGA Deckglas und lassen Sie die Substrate auf Raumtemperatur abkühlen. Sie PUA Form erst entfernen, wenn die Substrate wurdenabgekühlt. </li><li> Nachdem die Substrate ausreichend abgekühlt sind, vorsichtig abziehen der NP PUA Form und enthüllt die PLGA NP Substrat. Der NP PLGA Substrat sollte ein schillerndes Aussehen haben. </li><li> Platzieren abgeschlossen Proben Exsikkator zur Lagerung solange einen Monat. </li></ol> 4. Zellaussaat und Kultur HINWEIS: Dieses Protokoll beschreibt die Kultur der neonatalen Ratten-Kardiomyozyten (NRVMs) und H7 mit menschlichen embryonalen Stammzellen abgeleitete Kardiomyozyten (hES-CMs), aber andere Zellquellen verwendet werden. <ol><li> Bringen Sie die Deckgläser ANF (PU oder PLGA) zu einem 35-mm-Gewebekulturschale Polystyrol. Je 20 ul Norland Optical Adhesive (NOA83H) an den Boden der Schale und vorsichtig platzieren Sie die ANF Deckglas auf der NOA. Lassen Leim, um sich auszubreiten und decken gesamte Deckboden. Cure NOA indem Gericht für 10 min UV. </li><li> Sterilisieren ANF durch Spülen mit 2 ml 70% ige wässrige Ethanollösung für 5 Minuten, zweimal. E entfernenthanol durch Aspiration. Lassen ANF vollständig an der Luft trocknen für ca. 1 Stunde unter der UV-Sterilisationslampe (λ = 200-290 nm) in der biologischen Sicherheitsschrank. </li><li> Zelladhäsion wird durch Beschichten der in Fibronektin ANF Nacht verbessert. Verdünnen Fibronektin in DI Wasser auf 5 ug / ml. Pipette 2 ml Fibronektin-Lösung in die Schüssel. Legen im Brutschrank bei 37 ° C und 5% CO 2 über Nacht (mindestens 6 h). </li><li> Erhalten NRVMs, hES-CMs oder andere Herzzellen von Interesse nach den bisherigen Protokollen 22. </li><li> Zentrifuge Zellprobe bei 1.000 UpM für 3 Minuten, um die Zellen zu pelletieren. </li><li> Entfernen Sie vorsichtig den Überstand durch Aspiration. Stellen Sie sicher, dass das Pellet nicht zu stören. </li><li> Die Zellen in geeigneten Kulturmedien in einer Konzentration von 4,6 × 10 6 Zellen / ml. </li><li> Beim Pipettieren von 200 ul Zellsuspension auf sterilisiert ANF. Stellen Sie sicher, Zellsuspension auf dem Deckglas bleibt. </li><li> Platzieren Zellen bei 37 ° C und 5%CO 2 für 4 Stunden, damit Zellen auf ANF befestigen. </li><li> 2 ml warmes Kulturmedien unter den gleichen Bedingungen austeilen und ersetzen Zellen im Inkubator. </li><li> Nach 24 Stunden zu entfernen Medien und waschen mit 2 ml DPBS zweimal, um überschüssige Zellen zu entfernen. </li><li> 2 ml warmes Kulturmedien unter den gleichen Bedingungen austeilen und ersetzen Zellen im Inkubator. Kultur die Zellen bis zur Konfluenz. Ersetzen Medien jeden zweiten Tag. </li></ol>

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D., Bub, G., Lakireddy, P., El-Sherif, N. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Optical+imaging+of+arrhythmias+in+the+cardiomyocyte+monolayer.">Optical imaging of arrhythmias in the cardiomyocyte monolayer.</a> Heart Rhythm. 9, 2077-2082 (2012).</li><li>Kim, J., Hayward, R. C. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Mimicking+dynamic+in+vivo+environments+with+stimuli-responsive+materials+for+cell+culture.">Mimicking dynamic in vivo environments with stimuli-responsive materials for cell culture.</a> Trends in Biotechnology. 30, 426-439 (2012).</li><li>Henderson, D. J., Anderson, R. H. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=The+Development+and+Structure+of+the+Ventricles+in+the+Human+Heart.">The Development and Structure of the Ventricles in the Human Heart.</a> Pediatr Cardiol. 30, 588-596 (2009).</li><li>Badie, N., Bursac, N. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Novel+Micropatterned+Cardiac+Cell+Cultures+with+Realistic+Ventricular+Microstructure.">Novel Micropatterned Cardiac Cell Cultures with Realistic Ventricular Microstructure.</a> Biophysj. 96, 3873-3885 (2009).</li><li>Badrossamay, M. R., McIlwee, H. A., Goss, J. A., Parker, K. K. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Nanofiber+Assembly+by+Rotary+Jet-Spinning.">Nanofiber Assembly by Rotary Jet-Spinning.</a> Nano Lett. 10, 2257-2261 (2010).</li><li>Rao, 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=The+effect+of+microgrooved+culture+substrates+on+calcium+cycling+of+cardiac+myocytes+derived+from+human+induced+pluripotent+stem+cells.">The effect of microgrooved culture substrates on calcium cycling of cardiac myocytes derived from human induced pluripotent stem cells.</a> Biomaterials. 34, 2399-2411 (2013).</li></ol>

Authors have nothing to disclose.

Funktionell reifen Herzgewebe fehlen sowohl in vivo und in-vitro-Anwendungen von Herzgewebe Technik. Die hier beschriebenen Methoden sind CFL Nanofabrikation robuste Techniken zur Erreichung zelluläre Ausrichtung und Einfluss makroskopischen Gewebefunktion aufgrund der Skalierbarkeit des Systems. Große Flächen können einfach strukturiert und für die Zellkultur eingesetzt werden. Makroskopische zelluläre Ausrichtung ist notwendig, um biomimetische, funktionelles Gewebe, da sie sowohl mechanische u...

Kardiovaskuläre Erkrankungen sind die häufigste Ursache für Morbidität und Mortalität in der Welt und stellen eine gewichtige sozioökonomische Belastung für eine bereits angespannte globale Gesundheitssystem 1,17. Cardiac Tissue Engineering hat zwei unterschiedliche Ziele: (1), um beschädigte Herzmuskel nach einem ischämischen Kardiomyopathie Krankheit oder regenerieren oder (2), um eine High-Fidelity-Modell des Herzens für die in vitro Wirkstoff-Screening oder Krankheit Modellierung zu erstellen. Das Herz ist ein komplexes Organ, das ständig arbeiten muss, um Blut in den Körper zu versorgen. Dicht gepackte Schichtstrukturen von Kardiomyozyten und Stützgewebe werden in spiralförmigen Muster in der gesamten Herzwand 18,19 angeordnet. Das Herz wird auch elektromechanisch in einem sehr koordiniert, effizient Blut auswerfen, um den Körper 20 21 gekoppelt. Mehrere große Hürden angesprochen werden, jedoch vor komplizierten Design der Natur zuverlässig in vitro rekapituliert werden.Erste, wenn auch robust Kardiomyozytendifferenzierung Methoden weiterentwickelt werden 22, hpSC-CMs noch zeigen eher unreifen Phänotypen. Die elektromechanischen Eigenschaften und Morphologie am ehesten entsprechen fetalen Ebenen 23. Zweitens, wenn in der traditionellen Kulturbedingungen gehalten, sowohl Stammzellen gewonnenen Herzmuskelzellen und Grund nicht, in native, gewebeartigen Strukturen zusammenzubauen. Vielmehr Zellen werden zufällig ausgerichtet und nicht die gebändert stabförmige Aussehen der Erwachsenenherzmuskel 24 aufweisen. Die extrazelluläre Matrix (ECM), mit der Umgebung interagieren Zellen spielt eine wichtige Rolle bei zahlreichen zellulären Prozessen 11,13,25. Das ECM aus komplexen, gut definierte Molekular und topographischen Signale, die die Struktur und Funktion von Zellen 6,26 signifikant beeinflussen. Innerhalb des Herzens, sich eng an die zugrunde liegenden zellulären Ausrichtung Nanometerbereich ECM Fasern 2. Die Auswirkungen dieser nanotopographsche Signale auf Zell-und Gewebefunktion ist jedoch weit von vollständig verstanden. Voruntersuchungen der Nanometerskala zell Biomaterial Interaktion zeigen die potentielle Bedeutung und Auswirkung Submikrometerbereich topographischen Signale für Zellsignal 27, 28-30 Adhäsion, Wachstum 31 und Differenzierung 32,33. Jedoch aufgrund der Schwierigkeit bei der Entwicklung von reproduzierbaren und skalierbaren nanofabricated Substrate, wie Untersuchungen konnte die Multiskalen zellulären Effekte des Komplexes in vivo ECM-Umgebung zu reproduzieren. In diesem Protokoll wird ein einfacher und kostengünstiger Nanofabrikation Technik, um Zellkultur Gerüste imitiert nativen Herz ECM Faserausrichtung erzeugen beschrieben, so dass für eine breite Palette von neuen Untersuchungen von Kardiomyozyten-Biomaterial-Interaktionen. Verstehen, wie Kardiomyozyten interagieren mit dem nanoskaligen ECM-Umfeld konnte für die Fähigkeit, das Zellverhalten zu nativem Gewebe genauer nachahmen Funktionen steuern könnention. Darüber hinaus sind Zellmonoschichten eine vereinfachte Experimentalsystem im Vergleich zu 3D-Strukturen, aber immer noch komplexen mehrzelligen Verhalten für aufschlussreiche Untersuchungen und funktionelles Screening 2,34-36 aufweisen. Schließlich könnten solche Gerüste verwendet, um zelluläre Transplantatfunktion zu verbessern, wenn in das Herz für regenerative Zwecke 37 implantiert werden.

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capillary force lithography for cardiac tissue engineering

Herz-Kreislauf-Krankheit bleibt die führende Todesursache weltweit ein. Cardiac Tissue Engineering hält viel Versprechen, bahnbrechende medizinische Erkenntnisse mit den Zielen der Entwicklung von funktionalen Geweben für die Herzregeneration sowie in-vitro-Screening-Tests zu liefern. Allerdings hat die Fähigkeit, High-Fidelity-Modelle von Herzgewebe zu erzeugen schwierig erwiesen. Das Herz der extrazellulären Matrix (ECM) ist eine komplexe Struktur, bestehend aus beiden biochemischen und biomechanischen Signale im Bereich von der Mikro-bis Nanometerskala 2. Lokale mechanische Belastung und der Zell-ECM-Interaktionen wurden vor kurzem als wichtige Komponenten in der Herzgewebetechnik 3-5 anerkannt. Ein Großteil des Herz ECM von ausgerichteten Kollagenfasern mit Durchmessern im Nanomaßstab, die Gewebearchitektur und elektromechanische Kupplung 2 maßgeblich beeinflusst besteht. Leider haben nur wenige Methoden have in der Lage gewesen, die Organisation der ECM-Fasern bis in den Nanometerbereich zu imitieren. Jüngste Fortschritte in der Nano Techniken haben jedoch die Konstruktion und Herstellung von skalierbaren Gerüste, die die in vivo strukturelle Steifigkeit und Substrat Signale des ECM im Herzen 6-9 imitieren aktiviert. Hier die Entwicklung stellen wir zwei reproduzierbar, kostengünstig und skalierbar Nano Verfahren zur funktionellen Ausrichtung der Herzzellen mit dem biokompatiblen Polymer Poly (lactid-co-glycolid) (PLGA) 8 und ein Polyurethan (PU)-basierten Polymers. Diese anisotrop nanohergestellte Substrate (ANF) imitieren die zugrunde liegenden ECM gut organisiert, ausgerichtet Gewebe und kann verwendet werden, um die Rolle der Nanotopographie auf die Zellmorphologie und Funktion 14.10 untersuchen. Mit Hilfe eines nanostrukturierten (NP) Silizium-Master als Vorlage wird eine Polyurethan-Acrylat (PUA) Form hergestellt. Diese PUA Form wird dann verwendet, pattern die PU-oder PLGA Hydrogel über UV-unterstützte oder Lösungsmittel-vermittelte Kapillarkraft Lithographie (CFL), bzw. 15,16. Kurz gesagt, PU-oder Pre-PLGA-Polymer-Fall auf einem Deckglas und der PUA Form wird auf die Spitze verzichtet. Für UV-unterstützte CFL, die PU wird dann mit UV-Strahlung (λ = 250-400 nm) zur Härtung ausgesetzt wird. Für Lösungsmittel-vermittelte CFL wird die PLGA mit Hitze (120 ° C) und Druck (100 kPa) eingeprägt. Nach dem Aushärten wird die Form PUA abgezogen und hinterließ eine ANF für die Zellkultur. Primärzellen wie neugeborenen Ratten Kardiomyozyten, sowie menschliche pluripotente Stammzellen gewonnenen Herzmuskelzellen, kann auf der ANF 2 gehalten werden.

Fig. 1 ist eine schematische Übersicht des Herstellungsverfahrens für die beiden Herstellungsverfahren. Durch die Beugung des Lichts durch die nanoskaligen Topographie verursacht werden, sollten Nanostrukturierung in einem schillernden Oberfläche zu der ANF führen. Abbildung 2 zeigt diese irisierende Oberfläche auf einer gut strukturierten 25 mm NP-PU-Deckglas (2A) mit 800 nm Kamm und Nut Breite (Figur 2B). Die schillernde Erscheinung des ANF etwas...

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.

Kapillar-Force-Lithographie für Herz-Tissue Engineering

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 Aufrufe.  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.

Serie: 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.

Herstellung von biologisch abgeleiteten Injizierbare Materialien für Myocardial Tissue Engineering

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

Forschung

Biotechnik

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.

Elastomer-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 ...

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.

Plasma Lithography Oberflächenstrukturierung für die Erzeugung von Cell Networks

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.

Verkapselung von Kardiomyozyten in einer Fibrin-Hydrogel für 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 ...

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.

Postproduction Verarbeitung von Elektrogesponnene Fasern für 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 ...

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.

Entwurf eines Biaxiale Mechanische Belastung Bioreaktor für Tissue Engineering

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

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.

Tissue Engineering eines Menschen 3D<em&gt; In-vitro-</em&gt; Tumor-Testsystem

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.

Engineering 3D cellularized Kollagengele für Gefäßgeweberegeneration

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.

Herstellung von Thermoresponsive nanostrukturierte Oberflächen für Tissue Engineering

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.

Tissue Engineering von Intrinsic Vaskularisation in ein<em&gt; In Vivo</em&gt; 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 ...

Kapillar-Force-Lithographie für Herz-Tissue Engineering

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