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.

Todos los procedimientos se realizaron a temperatura ambiente (~ 23 ° C) a menos que se indique lo contrario. 1. Fabricación de Silicon Maestro <ol><li> Limpio oblea de silicio con etanol al 100% o xileno y se seca a O 2 / N 2 de gas. </li><li> Coloque oblea de silicio en Spin-revestidor a velocidades de rotación de 2.000-4.000 rpm para producir una película de espesor de 0,3-0,5 micras. </li><li> Patrón de la película fotorresistente con las dimensiones correctas mediante el uso de un sistema de fotolitografía </li><li> Sumergir completamente las obleas de silicio fotorresistentes recubiertos con dibujos en un volumen apropiado de solución fotosensible en desarrollo. </li><li> Enjuague las obleas de silicio recubiertas de resina fotosensible desarrollados con agua desionizada. </li><li> Para formar conjuntos de crestas escala inferior a la micra con paredes laterales casi verticales, profundidad de grabado iónico reactivo del silicio expuesta mediante un sistema de grabado. </li><li> Retire el fotoprotector restante mediante la colocación de la oblea de silicio en un sistema de Asher de plasma. </li><li> Cortar el silicio wAfers con un cortador de punta de diamante en los amos de silicio de tamaño adecuado para el moldeo réplica posterior. </li></ol> 2. Fabricación de PUA Molde de Silicon Maestro Nota: Volumen para ser añadido a maestro de silicio para la nanofabricación variará dependiendo de la zona del maestro nanoestructurada a ser replicado, así como la viscosidad de la solución de acrilato de poliuretano (PUA). <ol><li> Limpie la superficie de maestro de silicio con 100% de etanol o xileno y seco bajo O 2 / N 2 gas. </li><li> Coloque el lado patrón maestro de silicio en una placa de Petri. </li><li> Para un maestro de silicio con unos 2 cm x 2 cm de superficie, patrón de la pipeta 40 l de PUA a la superficie de diseño. </li><li> Coloque una hoja de 4 cm x 4 cm de poliéster (PET) de la película transparente sobre la PUA dispensado. </li><li> Presione hacia abajo sobre la lámina de PET y se extendió la PUA debajo de la hoja a través de la cara patrón usando un rodillo o una superficie plana filo (tal como una tarjeta) de modo que la totalidadpatrón está cubierto por el prepolímero PUA. </li><li> Coloque maestro de silicio, prepolímero, y PET de aproximadamente 10 cm por debajo de una luz de 20 vatios (115 V) UV (λ = 365 nm) durante 50 seg. Para ser efectiva, la longitud de onda de la luz UV puede ser en cualquier lugar entre 310-400 nm. La intensidad de la luz es de 10-15 mW / cm 2 en la superficie del sustrato. </li><li> Después del curado, retire la película de PET lentamente con fórceps. PUA debe adjuntar a la película de PET con un negativo de la maestra nanopattern silicio. </li><li> Curar nanopatterns PUA / PET bajo UV durante al menos 12 horas antes de su uso. La exposición excesiva no es un problema. </li><li> Para limpiar los maestros de silicio, coloque otra película de PET en la parte superior del maestro sin la adición de PUA y exponer a la luz UV (λ = 365 nm) durante 50 segundos y retirar la película de PET. Esto eliminará cualquier monómero sin reaccionar. </li><li> Enjuague maestro de silicio con 100% de etanol o xileno y seco bajo O 2 / N 2 gas. </li></ol> 3a. Polímero de poliuretano nanoestampación </p&gt; <ol><li> Preparar 25 mm de diámetro portaobjetos de vidrio circulares colocando en una cámara de tratamiento de ozono durante 10 min. </li><li> Coloque los portaobjetos de vidrio tratadas con ozono a un pequeño bloque de PDMS para un fácil manejo. </li><li> Aplique una capa delgada de la superficie de promotor de adhesión con la brocha a portaobjetos de vidrio. Aire portaobjetos de vidrio seco durante 30 min. </li><li> Coloque la placa de vidrio sobre un pedazo de papel de la impresora. </li><li> Gota dispensar 10 l de poliuretano (PU) de pre-polímero (NOA 76) hacia el centro de portaobjetos de vidrio. Asegúrese de que no hay burbujas después de la adición. </li><li> Molde Place PUA, patrón boca abajo, sobre el portaobjetos de vidrio. Dispersar la PU de manera uniforme en toda la superficie de la lámina de vidrio rodando un rodillo de cilindro de caucho a lo largo del molde PUA. Papel de impresora absorberá desbordamiento polímero. </li><li> Lámpara UV Watt. El tiempo de polimerización de la PU es dependiente de la energía de la fuente de UV. </li><li> Retire la muestra de la fuente de luz UV y pelar cuidadosamente el molde PUA de portaobjetos de vidrio con revestimiento de poliuretano. La polimerización de la PU se considera completa cuando el molde cáscaras PUA limpiamente lejos de la muestra y el portaobjetos de vidrio de la PU tiene un aspecto iridiscente. </li><li> Coloque las muestras de acabados en desecador para almacenarlo durante todo el tiempo como un mes. </li></ol> 3b. Nanoestampación poli (lactida-co-glicolida) Hidrogel <ol><li> Crear un molde de PDMS plana mezclando vigorosamente base de elastómero de silicona y silicona agente de curado de elastómero en una proporción de 10:1. </li><li> Pour PDMS mixtos precursor de solución en una placa de Petri de modo que el precursor de PDMS llega a 5 mm hasta el borde del plato (es decir, de modo que el PDMS es 5 mm de espesor). </li><li> Coloque el plato de Petri y PDMS precursor en un desecador durante 1 hora para desgasificar. </li><li> Mueva placa de Petri y PDMS precursor en un horno de 65 ° C durante al menos 2 horas para curar. </li><li> Después de PDMS es curado, utilizar una navaja de afeitar para cortar el PDMS planas en 3 cm x 3 cm secciones cuadradas. Estos serán los moldes de PDMS planas utilizadas más adelante en el proceso de modelado. </li><li> Clean diámetro 25 mm portaobjetos de vidrio circular por plsobresalir en alcohol isopropílico durante 30 minutos en un sonicador de agua. </li><li> Diapositivas lavados en seco de vidrio bajo O 2 / N 2 gas. </li><li> Gota dispensar 100 l de solución de PLGA (15% w / v de PLGA en cloroformo) en portaobjetos de vidrio. </li><li> Coloque un molde de PDMS plana en la parte superior de la dispensado de PLGA para absorber el disolvente y obtener una superficie plana de PLGA. Aplique una ligera presión (~ 10 kPa) mediante la colocación de un peso de 200 g en la parte superior de los PDMS durante 5 min. </li><li> Lentamente retire del molde de PDMS plana y colocar el vidrio de cubierta en una placa precalentado (120 ° C) durante 5 min para eliminar el disolvente residual y aumentar la adhesión entre el PLGA y el vidrio de cubierta. </li><li> Coloque el molde NP PUA en la parte superior de la plana de PLGA y aplicar presión constante (~ 100 kPa) y calor (120 ° C) mediante la colocación de un peso de 1 kg en la parte superior del molde PUA, mientras que en la placa de calentamiento durante 15 min. </li><li> Retire el peso de PLGA tapa deslizante y deje que el sustrato se enfríe a temperatura ambiente. No quitar el moho PUA hasta los sustratos han sidoenfriado. </li><li> Una vez que los sustratos se hayan enfriado lo suficiente, cuidado, retire el molde NP PUA, revelando el sustrato NP PLGA. El sustrato NP PLGA debe tener un aspecto iridiscente. </li><li> Coloque las muestras de acabados en desecador para almacenarlo durante todo el tiempo como un mes. </li></ol> 4. Siembra de células y Cultura NOTA: Este protocolo describe la cultura de los miocitos ventriculares de rata neonatal (NRVMs) y cardiomiocitos H7 humanos madre embrionarias de células derivadas de células madre (CMS), pero otras fuentes de células puede ser utilizado. <ol><li> Coloque el cubreobjetos ANF (PU o PLGA) a un 35 mm de cultivo de tejidos plato de poliestireno. Pipetear 20 l de Norland Optical Adhesive (NOA83H) a la parte inferior del plato y coloque suavemente el cubreobjetos ANF sobre el NOA. Deje que el pegamento hacia fuera y cubre toda la parte inferior cubreobjetos. Curar NOA exponiendo plato a los rayos UV durante 10 min. </li><li> Esterilizar ANF mediante aclarado con 2 ml de solución acuosa de etanol al 70% durante 5 minutos, dos veces. Retire eTHANOL por aspiración. Permitir ANF se seque por completo el aire durante aproximadamente 1 hora bajo la lámpara de esterilización UV (λ = 200-290 nm) en la cabina de seguridad biológica. </li><li> La adhesión celular se ve reforzada por el recubrimiento de la ANF en fibronectina durante la noche. Diluir la fibronectina en agua DI a 5 g / ml. Pipeta 2 ml de solución de fibronectina en el plato. Coloque en una incubadora a 37 ° C y 5% de CO 2 durante la noche (por lo menos 6 horas). </li><li> Obtener NRVMs, hESC-CM, u otras células cardiacas de interés de acuerdo con los protocolos anteriores 22. </li><li> Muestra de células Centrifugar a 1000 rpm durante 3 min para sedimentar las células. </li><li> Retire con cuidado el sobrenadante por aspiración. Asegúrese de no perturbar el sedimento. </li><li> Resuspender las células en medios de cultivo apropiados a una concentración de 4,6 x 10 6 células / ml. </li><li> Cuidadoso pipeteado 200 l de suspensión celular a ANF esterilizados. Asegúrese de que la suspensión de células se mantiene en el cubreobjetos. </li><li> Coloque las células en una incubadora a 37 ° C y 5%CO 2 durante 4 horas para permitir que las células se unan a la ANF. </li><li> Añadir 2 ml de medio de cultivo caliente al plato y reemplazar las células en una incubadora en las mismas condiciones. </li><li> Después de 24 horas, retirar el medio y lavar con 2 ml de DPBS dos veces para eliminar el exceso de células. </li><li> Añadir 2 ml de medio de cultivo caliente al plato y reemplazar las células en una incubadora en las mismas condiciones. Cultivar las células hasta confluencia. Vuelva a colocar los medios de comunicación todos los días. </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. 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Authors have nothing to disclose.

Tejidos cardíacos Funcionalmente maduros carecen de tanto in vivo como in vitro en aplicaciones de ingeniería de tejido cardíaco. Los métodos de nanofabricación CFL descritos aquí son técnicas robustas para lograr la alineación celular e influir en la función del tejido macroscópico debido a la escalabilidad del sistema. Grandes áreas pueden ser fácilmente modelados y se utilizan para el cultivo de células. Alineación celular macroscópico es esencial en la ingeniería de tejido cardiaco con el fi...

La enfermedad cardiovascular es la principal causa de morbilidad y mortalidad en el mundo y presenta una carga socioeconómica de peso en un 1,17 sistema mundial de salud ya tensas. Ingeniería de tejidos cardíacos tiene dos objetivos distintos: (1) para regenerar el miocardio dañado después de la enfermedad isquémica o miocardiopatía o (2) para crear un modelo de alta fidelidad del corazón para la detección in vitro de drogas o la modelización de enfermedades. El corazón es un órgano complejo que se debe trabajar constantemente para suministrar sangre al cuerpo. Estructuras laminares densamente poblado de cardiomiocitos y tejidos de soporte están dispuestos en patrones helicoidales a lo largo de la pared del corazón 18,19. El corazón también se acopla electromecánicamente 20 de una manera muy coordinada para expulsar de manera eficiente la sangre al cuerpo 21. Varios obstáculos importantes siguen pendientes de resolución, sin embargo, antes de que el intrincado diseño de la naturaleza de forma fiable puede ser recapitulado in vitro.En primer lugar, aunque los métodos de diferenciación de los cardiomiocitos robustos siguen elaborando 22, HPSC-CM, aún subsisten algunos fenotipos más inmaduros. Sus propiedades electromecánicas y morfología coincide más estrechamente los niveles fetales 23. En segundo lugar, cuando se mantienen en condiciones de cultivo tradicionales, tanto derivado de células madre y cardiomiocitos primarios logran ensamblarse en estructuras como tejidos nativos. Por el contrario, las células se vuelven orientados al azar y no presentan la apariencia en forma de barra con banda de miocardio adulto 24. La matriz (ECM) entorno extracelular con la que interactúan las células juega un papel importante en numerosos procesos celulares 11,13,25. El ECM consiste en señales moleculares y topográficas complejas, bien definidos que influyen significativamente en la estructura y función de las células 6,26. Dentro del corazón, alineación celular sigue de cerca las fibras de ECM escala nanométrica subyacente 2. El impacto de estos nanotopographseñales iCal en la célula y la función del tejido, sin embargo, está lejos de ser completamente entendido. Los estudios preliminares de la interacción célula-biomaterial escala nanométrica indican la importancia potencial y el impacto de sub-micrones señales topográficas para la señalización celular 27, la adhesión 28-30, un crecimiento del 31, y la diferenciación 32,33. Sin embargo, debido a la dificultad en el desarrollo de sustratos nanofabricados reproducibles y escalables, tales estudios podrían no reproducir los efectos celulares a escala múltiple del complejo ambiente de ECM en vivo. En este protocolo, una técnica de nanofabricación sencillo y rentable para producir andamios de cultivo de células que imitan la alineación de la fibra nativa ECM cardiaca se describe, lo que permite una amplia gama de nuevas investigaciones de las interacciones-cardiomiocitos biomaterial. La comprensión de cómo los cardiomiocitos interactúan con el ambiente de ECM nanoescala podría permitir la capacidad de controlar el comportamiento celular para reproducir con mayor fidelidad función tejido nativoción. Por otra parte, las monocapas celulares son un sistema experimental simplificado en comparación con las estructuras en 3D, pero todavía exhiben un comportamiento multicelular complejo para investigaciones profundas y selección funcional 2,34-36. Finalmente, tales andamios se podrían utilizar para mejorar la función celular del injerto cuando se implanta en el corazón para propósitos de regeneración 37.

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capillary force lithography for cardiac tissue engineering

La enfermedad cardiovascular sigue siendo la causa principal de muerte en todo el mundo 1. Ingeniería de tejidos cardíacos es muy prometedora para entregar los descubrimientos médicos innovadores con los objetivos de desarrollo de tejidos funcionales para la regeneración cardiaca, así como ensayos de cribado in vitro. Sin embargo, la capacidad de crear modelos de alta fidelidad de tejido del corazón ha demostrado ser difícil. Matriz extracelular del corazón (ECM) es una estructura compleja que consta de dos señales bioquímicas y biomecánicas que van desde la micro-a la escala del nanómetro 2. Local condiciones de carga mecánica y las interacciones célula-ECM han sido recientemente reconocida como componentes vitales en la ingeniería de tejido cardíaco 3-5. Una gran parte de la MEC cardiaca se compone de fibras de colágeno alineados con diámetros de nano-escala que influye significativamente en la arquitectura del tejido y el acoplamiento electromecánico 2. Por desgracia, algunos métodos VHAe sido capaz de imitar la organización de las fibras de ECM a la escala nanométrica. Los recientes avances en las técnicas de nanofabricación, sin embargo, han permitido el diseño y la fabricación de andamios escalables que imitan las señales in vivo de rigidez estructural y del sustrato en del ECM en el corazón de 6-9. Aquí se presenta el desarrollo de dos reproducible, procesos nanoestampación escalables para la alineación funcional de las células cardiacas utilizando el poli biocompatible de polímero (láctico-co-glicólico) (PLGA) 8 y una de poliuretano (PU) de polímero a base rentable y. Estos sustratos anisotrópicamente nanofabricados (ANF) imitan el ECM subyacente de tejidos, alineados bien organizadas y se puede utilizar para investigar el papel de nanotopografía sobre la morfología celular y la función 10-14. El uso de un nanoestructurada (NP) maestro de silicio como una plantilla, un acrilato de poliuretano (PUA) de molde se fabrica. Este molde PUA se utiliza entonces para PAttern hidrogel PU o PLGA mediante UV-asistida o fuerza capilar litografía-solvente mediada (CFL), respectivamente 15,16. Brevemente, PU o PLGA de pre-polímero es gota dispensado sobre un cubreobjetos de vidrio y el molde PUA se coloca en la parte superior. Para CFL UV asistida, la PU se expone entonces a la radiación UV (λ = 250-400 nm) para el curado. Para disolvente mediada CFL, el PLGA se graba en relieve el uso de calor (120 ° C) y presión (100 kPa). Después del curado, el molde PUA se despega, dejando atrás un ANF para el cultivo celular. Las células primarias, como los miocitos ventriculares de rata neonatal, así como los cardiomiocitos derivados de células madre pluripotentes humanas, se pueden mantener en la ANF 2.

La figura 1 es una vista general esquemática del proceso de producción para los dos métodos de fabricación. Debido a la difracción de la luz causada por la topografía a escala nanométrica, nanoestampación debe resultar en una superficie iridiscente a la ANF. Figura 2 representa esta superficie iridiscente en un 25 mm NP-PU cubreobjetos bien modelada (Figura 2A) con cresta 800 nm y la ranura ancho (Figura 2B). El aspecto iridiscente de la ANF var...

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.

Capilar Fuerza Litografía Cardiaca de Ingeniería de Tejidos

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

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12.3K vistas.  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.

Video: 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.

Fabricación de materiales de origen biológico inyectable para ingeniería de tejidos del miocardio

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

Investigación

Bioingeniería

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.

PGS elastomérico andamios en ingeniería de tejidos arteriales

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.

La litografía de plasma superficial Patrones de Creación de redes de telefonía

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.

La encapsulación de los cardiomiocitos en un hidrogel de fibrina para ingeniería de tejidos cardíacos

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.

Postproducción de procesamiento de las Fibras electrospun de Ingeniería de Tejidos

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.

Diseño de un Carga mecánica Biaxial biorreactor para Ingeniería de Tejidos

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.

Ingeniería de tejidos de un humano 3D<em&gt; In vitro</em&gt; Sistema de Prueba de Tumores

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.

Ingeniería 3D Cellularized colágeno Geles para Vascular Regeneración Tisular

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.

Preparación de Superficies de respuesta térmica nanoestructurados para la Ingeniería de Tejidos

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.

Ingeniería de tejidos por intrínseca Vascularización en una<em&gt; En Vivo</em&gt; Cámara Ingeniería de Tejidos

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

Capilar Fuerza Litografía Cardiaca de Ingeniería de Tejidos

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